CRANFIELD UNIVERSITY D J HOWORTH FEASIBILITY STUDY FOR A MICROMACHINED GLAUCOMA DRAINAGE DEVICE SCHOOL OF INDUSTRIAL AND MANUFACTURING SCIENCE

Transcription

1 CRANFIELD UNIVERSITY D J HOWORTH FEASIBILITY STUDY FOR A MICROMACHINED GLAUCOMA DRAINAGE DEVICE SCHOOL OF INDUSTRIAL AND MANUFACTURING SCIENCE MSC THESIS

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3 CRANFIELD UNIVERSITY SCHOOL OF INDUSTRIAL AND MANUFACTURING SCIENCE MSc Thesis Academic Year D J Howorth Feasibility Study for a Micromachined Glaucoma Drainage Device Supervisor: D Allen 13 September 2002 This thesis is submitted in partial fulfilment of the requirements for the degree of Master of Science Cranfield University All rights reserved. No part of this publication may be reproduced without the written permission of the copyright owner.

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5 ABSTRACT Current treatments for glaucoma are not fully satisfactory. This thesis examines the possibility of improving one of those treatments the use of a drain in the eye. The goal of the thesis is to increase the likelihood that improved treatments for glaucoma become available at an early date by the use of microsystems engineering. Glaucoma is a serious and common disease of the eye, most often caused by raised pressure within the eye. It is usually treated by drugs or surgery but sometimes these treatments cannot be used and various artificial drainage implants are then used to relieve the overpressure, but with limited success. This thesis is in the nature of a feasibility study, exploring the design space for glaucoma drainage devices in order to identify opportunities to advance the state of the art by applying microsystems technology. The document contains a summary of the medical background and reviews of intraocular pressure sensors and current drainage devices. It considers the requirements and issues facing an improved device and proposes the use of a micromachined, pressure-controlled active valve within a drainage device. It examines in detail possible actuators for the valve and the means to power it.

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7 TABLE OF CONTENTS 1 INTRODUCTION Structure of Thesis MEDICAL BACKGROUND Anatomy of the eye...3 Overall structure...3 Aqueous humour...5 Intraocular lens Glaucoma Intraocular Pressure...7 Hypotony Conventional Tools and Therapies...9 Pressure measurement techniques and problems...9 Applanation tonometry...9 Indentation tonometry and tonography...10 Manometry...10 Current treatments for glaucoma PRESSURE SENSORS Ambient Pressure...13 Relative and absolute pressure sensors...13 Change in atmospheric pressure with altitude...14 Medical significance...15 Ocular example Intraocular Pressure Sensors Summary List Early Projects...18 Collins...18 Cooper and Beale Sweden - Uppsala...18 Who...19 What...19 Results and analysis...19 thesis_djh_gdd.doc i

8 Table of Contents 3.5 Germany Aachen, Köln, Duisberg...20 Who...20 Publications...21 What...21 Results and analysis Germany Bremen, Hamburg, Zürich et al...23 What...24 Results and analysis Germany Karlsruhe...25 What...25 Results and analysis Belgium KU Leuven...26 What...26 Results and analysis France Tronic s...28 Who...28 What...29 Results and analysis Other Work...29 Netherlands University of Twente...29 Sweden Umeå University...30 US University of Minnesota Conclusions DRAINAGE IMPLANTS...31 Sources...31 History Survey of Current Types...32 Bleb...32 Molteno...33 Schocket...33 White...34 Krupin...34 Baerveldt...34 Ahmed...34 ii thesis_djh_gdd.doc

9 Table of Contents OptiMed Problems...35 Hypotony...36 Bleb overgrowth...36 Tube fit...36 Tube blockage...36 Valves not closing New Devices...37 AGFID...37 Porous materials...38 Schlemm s canal implants Patents DESIGN CONSIDERATIONS System Design...41 Is the pressure sensor integrated?...43 Is the active valve integrated? General Requirements...44 Size...44 Biocompatibility...45 Blockage...45 Cost...45 Reliability...46 Pressure...46 Temperature...47 Drain and valve...47 Flexibility (ability to change parameters)...47 Self-test...48 Regulatory...48 Orientation...48 Noise...49 Power supply availability...49 Bilateral implantation Viscosity of aqueous Other Issues...50 thesis_djh_gdd.doc iii

10 Table of Contents Entry port...50 Transport...51 Exit port...51 Soft lithography Biocompatibility and Materials...53 Patient health...53 Survival of the device...54 Encapsulation...54 Sterilisation...54 Biomaterials...55 PDMS...55 Polyimide...55 PMMA...56 PTFE...56 Apatite...56 Coatings VALVE AND ACTUATOR DESIGN Actuation Principles...60 Piezoelectric actuators...60 Heated actuators...61 Direct liquid drive actuators...61 Magnetic actuators...61 Summary Electrochemical Bubble Actuators...63 Principle...63 Advantages...64 Challenges...64 Neagu...65 Papavasiliou...67 Other aspects...70 Summary Electrostatic Actuators...72 Principle...72 Fabrication...74 SUMMiT...74 iv thesis_djh_gdd.doc

11 Table of Contents Out of plane actuation...75 Power supply...77 Force...77 Latching mechanisms...80 Interface circuitry Valve...83 Membrane POWER SUPPLY DESIGN Power Inductive loops...86 Donaldson and Perkins...87 Van Schuylenbergh et al...87 Neagu et al...87 Ullerich et al...88 Eggers et al...89 Permeable Cores...90 Other points Coils for Power Transfer...91 Coil position...91 Conventional coil technology...92 NMR microcoils...92 Other manufacturing techniques Other techniques for Power Transmission...93 Optical...94 RF...94 Biochemical Voltage and Power Considerations...95 Voltage Multipliers...96 High voltage CMOS...99 Micromechanical switches Communication Communication from external world into device Communication from device to external world Summary thesis_djh_gdd.doc v

12 Table of Contents 8 CONCLUSIONS Device Overview Research Topics Viscosity of aqueous Electrostatic power Electrochemical microcells Project Organisation Summary ACKNOWLEDGMENTS REFERENCES A DESCRIPTION OF THE EYE A.1 Introduction A.2 Gross anatomy and physiology of the eye A.3 External features of the globe Scleral/corneal layer Vascular/choroidal layer Retinal layer A.4 Internal features of the globe The lens The aqueous humour The vitreous humour B UNITS B.1 Pressure B.2 Flow Rates and Volumes C GLOSSARY C.1 Materials and Micromachining C.2 Electronics C.3 Ocular vi thesis_djh_gdd.doc

13 Table of Figures TABLE OF FIGURES Figure 1 The eyeball...4 Figure 2 Influences on intraocular pressure...7 Figure 3 Goldmann tonometer...9 Figure 4 Relative and absolute pressure sensors...14 Figure 5 Diagram of forces acting on the sclera...16 Figure 6 Passive resonant pressure sensor...19 Figure 7 Difference between the sensor signal and a reference barometer...20 Figure 8 Overall concept of OPHTHAL system...21 Figure 9 Block diagram of OPHTHAL components...22 Figure 10 The foldable pressure sensor Project IODS Figure 11 Prototype intraocular pressure sensor...24 Figure 12 Resonance frequency of a pig eye...25 Figure 13 Position of Leuven sensor...26 Figure 14 Passive harmonic telemetry...27 Figure 15 Passive resonant distributed LC pressure sensor...28 Figure 16 General arrangement of Tronic s sensor...28 Figure 17 Drainage device on the surface of the eyeball...32 Figure 18 Molteno implant...33 Figure 19 Baerveldt implant...34 Figure 20 Ahmed implant...35 Figure 21 OptiMed implant...35 Figure 22 AGFID prototype...37 Figure 23 AGFID drawings...38 Figure 24 Drainage device with porous anchorage...38 Figure 25 Schlemm s canal drainage device...40 Figure 26 System design for active pressure-controlled GDD...42 Figure 27 Simple valve cast in PDMS...53 Figure 28 The effect of coating PMMA...57 Figure 29 Principle of piezoelectric actuator...60 Figure 30 Principle of magnetic actuator...62 Figure 31 Hydrolysis cell...63 Figure 32 Cross section of electrochemically actuated valve...65 Figure 33 Construction of Neagu s actuator...66 Figure 34 Cross-section of Papavasiliou s valve...68 Figure 35 Electrochemical actuator by Papavasiliou...69 thesis_djh_gdd.doc vii

14 Table of Figures Figure 36 Principle of electrostatic actuator...73 Figure 37 SUMMiT process...74 Figure 38 Sandia mirror deflection system...76 Figure 39 Buckling beam valve actuator...76 Figure 40 Electrostatic actuator...78 Figure 41 Preliminary sketch of electrostatically actuated valve...79 Figure 42 Modular geared transmission...80 Figure 43 Worm gear...81 Figure 44 Sketch of pawl and rack...81 Figure 45 Principle of buckling beam...82 Figure 46 Microswitch using buckling beams...82 Figure 47 Principle of induction coils...87 Figure 48 Fabrication process for flexible planar coil...89 Figure 49 Planar microcoil with permeable layer...90 Figure 50 Villard cascade voltage multiplier...96 Figure 51 Parallel voltage multiplier...97 Figure 52 Four-stage Dickson charge pump...98 Figure 53 Charge pump by Shin...98 Figure 54 Micromechanical switch for charge pump Figure 55 Low-voltage RF MEMS switch Figure 56 Overview of active pressure-controlled GDD Figure 57 The structure of the eye viii thesis_djh_gdd.doc

15 1 INTRODUCTION Glaucoma is a serious disease of the eye, often associated with raised pressure within the eye. The disease is often undiagnosed until it has progressed to the extent of causing significant damage to the vision and if left untreated, it causes blindness. It is also a common disease, affecting up to 2% of the general population and over 10% of specific groups such as black men and Hispanic women over 80 in the U.S. It can be controlled by three types of treatment but none of these is entirely satisfactory and one type glaucoma drainage devices has a particularly mixed record. Advances in technology and understanding over the last decade have led people to consider whether it is possible to improve these devices through new designs that incorporate novel materials and fabrication technologies. Such a cycle of development has already taken place in the related field of sensors to measure the pressure within the eye and new micromachined sensors are about to enter clinical trials. These and other components that might be used to make an improved glaucoma drainage device are surveyed in this thesis and a design is proposed for a new device. The proposed drainage device incorporates an active, pressure-controlled valve to regulate the flow of liquid out of the eye. It is important to remember that the final goal is to put the device into production for clinical use. Thus safety, tried and tested, adequate, easy to produce are all terms that outweigh novelty, improved performance, et cetera when considering designs. There are two principal audiences for this document. Some have specialist medical knowledge, of ophthalmology and glaucoma in particular, while others are experts in the field of microsystems and especially in materials, processes and design. These two fields of study are separate with little commonality and consequently the document includes material intended to enable both sets of reader to benefit from all chapters. 1.1 Structure of Thesis The next chapter describes the basic anatomy of the eye, the condition of glaucoma and the mechanism that sustains the pressure within the eye. Finally, it summarises some of the medical devices used to measure pressure and the methods of treatment available for glaucoma. thesis_djh_gdd.doc

16 Introduction Chapter 3 reviews the various projects that are building pressure sensors that can be implanted in the eye. It also examines the effects of changes in atmospheric pressure and the implications for drainage device design. Attention then turns to the existing drainage devices that are currently used in clinical practice, together with the problems that can arise with these devices and some new developments that are already in progress. Design considerations are the subject of chapter 5, which covers a range of topics that affect the design of an improved drainage device. The following two chapters deal with specific aspects of the design of a drainage device that uses an embedded actively-powered, pressure-controlled valve to regulate the flow rate. Chapter 6 explores the design of actuators to open and close the valve, whilst chapter 7 investigates how these actuators can be powered. These chapters are followed by the conclusions. There are three appendices that provide more information about the eye, about the units used to measure pressure and about the terms used in this document. Readers are advised to read Appendix B Units now thesis_djh_gdd.doc

17 2 MEDICAL BACKGROUND This chapter provides a brief, simplified introduction to the medical knowledge necessary to understand the technical issues of the drainage device and pressure sensor. It does not attempt to provide a complete background but concentrates on those details that are important elsewhere in this thesis. The next section introduces the anatomy of the eye. A more comprehensive summary, from (Lloyd 2001), is reproduced as Appendix A and details can be found in the main sources used to prepare this section, (Snell 1998) and (Kanski 1999). After that, a similar introduction is given to the medical conditions known as glaucoma and to the intraocular pressure that is the underlying cause of most cases. Finally there is a brief description of the range of treatments currently used for glaucoma. Throughout, the emphasis is on the liquid flow, pressure and other aspects that are important in the design of drainage devices. 2.1 Anatomy of the eye A horizontal section through the eyeball at the level of the optic nerve is shown in Figure 1. Many of the labelled terms are described in succeeding sections. Overall structure The eyeball is roughly spherical and about 24 mm in diameter. It is contained by the sclera, which is mostly opaque and white but blends into the transparent cornea at the front in order to admit light to the eye. The retina lines the inside rear half of the eyeball, registering the light falling on it and sending signals to the brain via the optic nerve. The bulk of the eye is filled and supported by a clear, jelly-like substance called the vitreous body. Near the front of the eye is the elastic lens which provides adjustment of focus. It is suspended within a capsular bag from the ciliary body, which contains the muscles that change the focal length of the lens by pulling it into a flatter shape. The small volume in front of the lens is of most interest in this thesis. The volume is divided into two by the iris, which controls the aperture of the lens and so the amount thesis_djh_gdd.doc

18 Medical Background of light striking the retina; the pupil is the hole in the centre of the iris through which light passes. The volume between the iris and lens is the posterior chamber (60 µl) and that between the iris and the cornea is the anterior chamber (250 µl). Both chambers are filled with a clear liquid known as aqueous humour or simply aqueous. Figure 1 The eyeball A horizontal section. The front of the eye is at the top. (Snell 1998) Also of interest is the space outside the sclera. The eye is encapsulated by a membrane known as Tenon s capsule over the rear portion that is not visible. The visible portion of the eye is covered by another membrane the conjunctiva which also lines the inside of the eyelids. These two membranes isolate the eye from the outside and their integrity is vital to prevent infections entering the eye. The interface between these membranes and the sclera forms a region that is used by many of the glaucoma drainage devices that are discussed later thesis_djh_gdd.doc

19 Medical Background Aqueous humour The aqueous is the clear liquid that fills the anterior and posterior chambers of the eye. It has very similar properties to water, which forms over 98% of it. Aqueous fulfils two main roles apart from forming part of the optical path within the eye: It supplies the oxygen and other metabolic demands of the lens and some of the cornea and it maintains the pressure within the eye. Aqueous is continuously formed by the ciliary body in the posterior chamber by secretion from the blood vessels there. It flows around the iris and lens into the anterior chamber, nourishing the tissues as it passes. It then leaves the eyeball through the trabecular meshwork, a sieve-like structure situated at the corner of the iris and the wall of the eye (the corner is known as the iridocorneal angle). Most of the aqueous filters through the trabecular meshwork into Schlemm s canal, a small channel that drains into the veins. A smaller portion rejoins the circulation after passing through the ciliary body and eventually through the sclera (the uveoscleral route). Other methods of formation and drainage are ignored here. The balance between the rate of formation and the rate of drainage of aqueous controls the pressure in the eye. The rate of formation is typically relatively constant at µl/min whilst the outflow varies, in particular rising with the pressure itself. This subject is explored in the section on Intraocular Pressure. Blood cells, proteins and other particles are found in the aqueous sometimes, especially after an operation such as that to implant a device. The viscosity of aqueous is uncertain; this is an important factor in governing flow rates through drainage devices and is discussed further in chapter 5. Intraocular lens The intraocular lens is of interest only through the sad fact that it is routinely replaced by an artificial lens in operations to cure cataracts a condition in which the lens has become obscured. Another unhappy but useful statistic is an 11% coincidence of cataract and glaucoma in the same patient (Eggers 2000b). The lens thus provides one possible place to position an embedded pressure sensor such as those discussed in the next chapter. its diameter varies between 350 and 500 µm thesis_djh_gdd.doc

20 Medical Background 2.2 Glaucoma Glaucoma is the name for a variety of serious conditions that can lead to blindness but no single definition encompasses all forms. Optic nerve damage is the cause of the blindness and it often leads first to a loss of visual field. The damage is most often caused by raised pressure and this rise is frequently caused by the blockage of the drainage paths for aqueous through the trabecular meshwork. However, it is possible for damage to occur with no increase in pressure (over 25% of cases) and conversely some individuals have a raised pressure without suffering damage. Pain is sometimes a symptom but frequently the damage is done before any symptoms appear. This section does not attempt to describe the various types of glaucoma, instead it merely seeks to present specific aspects that are salient to the thesis as a whole. A description of the glaucomas can be found in chapter 6 of (Kanski 1999) and the multi-volume work edited by Ritch, Shields and Krupin gives comprehensive coverage of the subject (Ritch 1996). These sources were used to prepare this section. Glaucoma is responsible for the loss of vision in 15% of the registered blind people in Germany. The disease does not usually cause pain or any other symptoms except loss of vision in the later stages. Therefore, glaucoma has typically already destroyed more than 80% of the nerve fibres before it is detected. (Walter 2000b) The two most important points are that [1] knowledge of the pressure within the eye is vitally important in the diagnosis and treatment of glaucoma and [2] control of the pressure, particularly reduction of raised pressure (hypertension), is key to controlling most forms of the disease. Many forms of glaucoma are caused by a reduction in the outflow of aqueous, leading to an increase in pressure. Thus another important point is that [3] increasing the rate of outflow is frequently desirable. The reduction in outflow can be from a variety of causes including a change in geometry that prevents aqueous reaching the trabecular meshwork (the iridocorneal angle is said to be a closed angle), overgrowth of the trabecular meshwork or its obstruction by particulate matter. There is also much information on the Web at sites such as the Glaucoma Research Foundation, thesis_djh_gdd.doc

21 Medical Background The possible presence of particles suspended in the aqueous is worth noting separately. These particles can include red blood cells, haemolysed red blood cells (so called ghost cells), proteins and pigment particles among others. Additional debris is also found in the aqueous after surgery. These particles are an important consideration in the design of a drainage device as well as in the etiology of glaucoma. Glaucoma is a common disease. According to a recent report, it affects 2.2 million people in the USA, 1.9% of the population. Rates rise to over 10% for some groups of people aged 80 or more (PBA 2002). 2.3 Intraocular Pressure The pressure within the eye is referred to as the intraocular pressure or IOP. Figure 2 is a general diagram of the many influences on the IOP. The diagram is from (Collins 1980), which with chapter 14 of (Ritch 1996), by Kaufman, provides the material in this section. Figure 2 Influences on intraocular pressure The arrows denote the direction of influence; dashed lines indicate a negative relationship (e.g. an increase in venous pressure causes a decrease in blood flow) (Collins 1980) The main relationships can be understood by taking the most significant factors first. As previously described, the IOP depends on the balance between the rate of formation of aqueous humour and its rate of outflow at steady state these must be equal. The IOP also depends on the pressure in the veins around the eye (the thesis_djh_gdd.doc

22 Medical Background episcleral venous pressure). These lead to the following expression for the rate of outflow through the trabecular meshwork: Q t = C f ( v IOP P ) where Q t is the rate of aqueous outflow (typically between µl / min) C f is the aqueous outflow facility (typically 0.2 µl / min / mmhg ) P v is the episcleral venous pressure (typically 9 mmhg) This equation, first stated by Friedenwald, is consistent with Poiseuille s formula for flow through a tube. The uveoscleral outflow, Q u, is believed to be independent of pressure for eyes that are not hypotonic, though estimates of the proportion of the total flow that it represents vary from 5% to 30%. The rate of formation of aqueous is inversely dependent on the IOP, though less sensitively than the outflow: Q p = C p ( P IOP) c where Q p is the rate of aqueous production (typically µl / min) C p is the aqueous production facility (approximately µl / min / mmhg) P c is a cutoff pressure at which aqueous formation is completely inhibited (approximately 90 mmhg) The normal range of IOP is from 10 to 22 mmhg, with a typical value of 16 mmhg.. Deep respiration can produce a rise and fall in the IOP of 5 mmhg and the pulse produces changes of up to 2 mmhg. In addition, IOP varies throughout the day with a typical range of 5 mmhg and is usually highest in the morning. Hypotony Medical problems arise when the IOP is too low as well as when it is too high. This condition, known as hypotony, is usually defined as an IOP of less than 6 mmhg. It is important in the treatment of glaucoma and in the design of glaucoma drainage devices mainly as an undesirable side-effect of surgery or the implantation of a drainage device and is discussed further in section mmhg = 267 Pa; 5 mmhg = 667 Pa; 6 mmhg = 800 Pa; 9 mmhg = 1200 Pa; 10 mmhg = 1333 Pa; 16 mmhg = 2100 Pa; 22 mmhg = 2900 Pa; 90 mmhg = Pa; µl/min/mmhg = µl/min/pa; 0.2 µl/min/mmhg = µl/min/pa thesis_djh_gdd.doc

23 Medical Background 2.4 Conventional Tools and Therapies This section gives a brief description of relevant tools and treatments for glaucoma. The most important of these from the viewpoint of this thesis are expanded in subsequent chapters. Pressure measurement techniques and problems Current techniques for measuring IOP suffer from a number of problems. - All the techniques must usually be performed by medically qualified staff and may involve the use of drugs (anaesthetics and dyes) and costly equipment, which consequently requires a visit to a hospital. - Because of this inconvenience, a limited number of samples are available, especially at night, which can lead to undesirable extrapolation when making a diagnosis. - Again because of the inconvenience, patients can become agitated with the result that the value to be measured changes. The measurement affects the measurand a problem that medical science shares with quantum science! - There is a problem of limited accuracy. The measurement is taken from outside the eye in almost all techniques, even though the pressure to be measured is inside the eye. This is necessary to avoid the risks and unpleasantness of the surgical procedure that is necessary to measure the pressure within the eye, but it leads to a significant inaccuracy due to the unknown stiffness of the wall of the eye (usually the cornea). The stiffness often varies in diseased eyes. Measurement of IOP is called tonometry and the instruments that are used to perform it are called tonometers. Applanation tonometry The most accurate method of measuring IOP in regular use is applanation tonometry using a Goldmann tonometer, right. Applanation refers to the operating principle of the device, in which the force needed to flatten a small circle Figure 3 Goldmann tonometer thesis_djh_gdd.doc

24 Medical Background of the cornea is measured. Even though it is the gold standard technique, it suffers from all four of the problems described above. The air-puff tonometer used by high-street optometrists is also an applanation tonometer, but it sacrifices accuracy for convenience of use. Indentation tonometry and tonography With the indentation tonometer, the indentation in the cornea produced by a constant force is measured and used to calculate the IOP. By taking continuous readings over a period of a few minutes, it is also possible to estimate the rate of outflow of aqueous. This process is called tonography. Manometry Manometry is the most accurate method of measuring IOP but it is invasive so is almost never used. A tube is inserted into the eye so that the pressure within the eye can be measured directly with a U-tube manometer. Historically it has only been used during experiments but not for clinical practice. Two researchers in the mid-90s implanted catheters and remote telemetered pressure sensors in rabbits, obtaining new information about pressure variations due to natural cycles and to handling (McLaren 1996), (Schnell 1996). Recently, an adaptation using a fine needle and an electronic pressure sensor has been introduced into clinical practice in difficult cases (Marx 1999). These results perhaps give the clearest indication of the value of an implanted pressure sensor. Current treatments for glaucoma There are presently three principal therapies for raised IOP: medical treatment with drugs, glaucoma filtration surgery (GFS) and implantation of a glaucoma drainage device (GDD) to assist drainage. Which of these is given to each patient depends on a number of factors including the risks associated with the treatment, the likelihood of its success and the clinical presentation of the individual patient thesis_djh_gdd.doc

25 Medical Background Medical treatment is usually given first since it has a good success rate and requires no surgical intervention with its associated risks. Drugs are available that reduce the rate at which aqueous forms and other drugs increase the rate of drainage through the trabecular meshwork. However, these drugs do not succeed in all cases and they also have side-effects that can be severe some are known to be carcinogenic so there is a need for alternative treatments. Surgery is the treatment of second choice and is becoming increasingly popular. Traditionally, this involves cutting a flap in the sclera, creating a new path through which the aqueous can drain into a pocket between the sclera and the conjunctiva (trabeculectomy). A newer alternative is to use a laser on up to 100 very small spots in the trabecular meshwork from the inside, which improves flow through the meshwork (trabeculoplasty). Such wounds naturally heal, leading to a need to repeat the surgery though this can be delayed by using drugs in conjunction with the surgery. Glaucoma filtration surgery is nevertheless considered a successful and very useful treatment. Glaucoma drainage devices (GDD) have a chequered history. There are many clinical complications which can lead to failure of the implant and this has led to GDDs being used as a treatment of last resort, when other alternatives have failed. This in itself makes the success rate lower; for example, many of the eyes are already scarred through repeated GFS and the patients sometimes die from an underlying disease that is causing the glaucoma (usually diabetes mellitus), leading to poor statistics for the GDDs. More detail about GDDs is provided in chapter 4. Improvements to the design of GDDs that lead to increased success rates might also enable it to be substituted for medical treatment or surgery in more cases. thesis_djh_gdd.doc

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27 3 PRESSURE SENSORS The pressure sensor is an essential part of a glaucoma drainage device incorporating an active valve. This chapter reviews pressure sensors that have been designed for implantation in the eye for diagnostic purposes and concludes that choosing one of these is a satisfactory basis for the drainage device design. The power supplies developed for these sensors are also relevant for the drainage device and are considered in a later chapter. Before reviewing the devices themselves, there is an explanation of the relevance of atmospheric pressure to the design range of these devices, which is also considered later. 3.1 Ambient Pressure Variations in atmospheric pressure place technical requirements on medical pressure sensors. Ambient air pressure varies somewhat with the weather and more significantly with changes in elevation. It also varies in two man-made situations flying and diving. The body s natural responses are designed to compensate for these changes and are supplemented by pressurised air supplies and other technology when necessary, but the nature of pressure sensors and their environment in the body requires that further special measures are taken. These factors are explored below. Relative and absolute pressure sensors Most pressure microsensors in clinical use the principle of measuring the deflection of a membrane caused by the difference in pressure between its two sides (other types of sensor change the details but not the nature of this discussion). Sometimes, as when the sensor is on the end of a catheter, it is possible to transmit the pressure to be measured through a fluid to an external point where it can be applied to one side of a membrane and atmospheric pressure can be applied to the other (equivalently, atmospheric pressure can be transmitted through a fluid to the point of interest and applied to a membrane there). Such a sensor is known as a relative pressure sensor and is said to measure gauge pressure; the principle is sketched in Figure 4. However, in the case of a totally enclosed sensor, such as an implant in the eye, there is no way to transmit the external pressure to the membrane. In these cases, the sensor thesis_djh_gdd.doc

28 Pressure Sensors pressure to be measured membrane pressure to be measured ambient pressure reference absolute pressure reference Relative Sensor Absolute Sensor A B Figure 4 Relative and absolute pressure sensors is designed such that one side of the membrane is exposed to a sealed cavity with a vacuum inside it (or sometimes a low-pressure gas) and the sensor measures the difference between the pressure of interest and the sealed-in reference pressure. Such sensors are called absolute pressure sensors and their readings do not change when the atmospheric pressure changes. Change in atmospheric pressure with altitude Atmospheric pressure descreases with increasing altitude. For the purposes of the discussion here, the rate of change can be approximated as 12.5 Pa for every metre of height (or 1 mmhg for every 11 m). At a height of 3000 m the pressure has reduced by 37.5 kpa (~280 mmhg) and the partial pressure of oxygen is becoming insufficient to support life without artifical aids. Normal passenger aircraft operate much higher than this, so have to be pressurized to sustain the passengers and crew. However, pressurization is expensive in terms of the weight of the fuselage and compression equipment, so a compromise is adopted. Modern passenger jets maintain the cabin pressure at an altitude of about 2000 m that is, reduced by 25 kpa or 190 mm Hg. This lower pressure also reduces the dangers from a sudden explosive decompression. Military aircraft operate with much lower pressures for performance reasons. Diving presents the opposite situation. Pressure increases by one atmosphere (approx 100 kpa or 760 mmhg) for every 10 m of depth. Floating on the surface face down with the head immersed may increase the pressure on the eyes by 2.5 kpa (19 mmhg) thesis_djh_gdd.doc

29 Pressure Sensors These changes of pressure have particular significance in mountainous regions such as the Alps where patients must change altitude to reach medical facilities and also in the US where flying is the most practical method of reaching hospital for some patients. Medical significance The body s systems operate on pressure differences (relative pressures) between ambient and internal and cope well with changes in atmospheric pressure. The solid and liquid materials from which the bulk of the body is formed expand or contract slightly as the ambient pressure is reduced or increased respectively and this brings the relative pressures back to near their original values. Gas within the body changes volume much more significantly and pressures must be equalized. Well known examples are the adjustment of the mass of air in the inner passages of the ear when climbing hills or flying and the exhalation of air from the lungs of divers, which can be seriously damaged if the diver fails to exhale upon rising. Ocular example Suppose that we had implanted a GDD with a flow rate valve controlled by an implanted absolute pressure sensor and suppose the patient started to ascend a hill. As the atmospheric pressure decreases whilst the absolute pressure inside the eye remains constant, the stress on the sclera increases and it stretches slightly (see Figure 5). This causes the volume inside the sclera to increase a little and the absolute pressure inside the eye reduces in consequence. If the sclera were extremely elastic (i.e. Young s modulus very small) it would expand until the stress across it regained its original value, that is until the relative pressure (IOP) regained its original value; in reality the IOP increases slightly because of the increased tension in the sclera. So the tissues of the eye have adjusted themselves such that the IOP retains its reasonable clinical value. Unfortunately, the outcome with the GDD is not so happy. thesis_djh_gdd.doc

30 Pressure Sensors atmospheric pressure IOP tension in sclera Figure 5 Diagram of forces acting on the sclera The pressure sensor registers the drop in absolute pressure inside the eye and this is interpreted to mean that fluid is flowing out of the eye too rapidly. So the valve is closed, preventing drainage from the eye and leading to an increased volume of fluid within the eye, until it restores the original absolute pressure and the regulator starts to open the valve again. At this point, the IOP has been elevated. This effect could have clinical consequences very quickly. Walking up a hill of 100 m would cause an increase in IOP of 9 mmhg (1200 Pa), climbing a 900 m hill would increase the IOP to over 90 mmhg. This is high enough to stop production of aqueous and blood flow in the eye. Conversely, walking down a 200 m hill from the hospital would cause the GDD to open the valve until all pressure in the eye was lost. It is thus absolutely essential that an actively-controlled valve in a GDD is controlled by relative pressure values rather than absolute ones. This means that a pressure reading is needed from some point known to be close to or at atmospheric pressure, as well as one from inside the eye. The uveitic outflow has been ignored thesis_djh_gdd.doc

31 Pressure Sensors 3.2 Intraocular Pressure Sensors Summary List Date Author Sensor Type Communication Location Source 1967 Collins pressure sensor passive RF resonator coils soft contact lens (Collins 1967) 1974 Greene and Gilman applanation tonometer resonant capacitor-coil corneal surface (Greene 1974) 6 mm dia, 3 mm thick (moveable ferrite core) 1977 Cooper and Beale commercial strain gauge wire link soft contact lens * 1983 Bramm bubble tonometer passive RF resonator coils implanted in vitreous (Bramm 1983) or anterior chambers 1990 Bäcklund et al capacitive, silicon passive coil / capacitor resonator synthetic lens implant * 1995 Puers capacitive resonator coil synthetic (PMMA) lens implant * 1996 Van Schuylenbergh capacitive passive coil / capacitor resonator choroid surface(van Schuylenbergh 1996a) 1996 McLaren et al commercial sensor battery-powered radio transmitter subcutaneous implant on neck * with catheter into eye 2000 Schnakenberg et al capacitive, silicon active RF induction link synthetic lens implant * 2000 Eggers et al capacitive, silicon active RF induction link synthetic lens implant (Eggers 2000b) 2000 Renard et al capacitive, SOI active RF induction link generic (Renard 2000) 2001 Rizq et al piezoresistive n/a choroid surface (Rizq 2001) * details from (Schnakenberg 2000) thesis_djh_gdd.doc

32 Pressure Sensors 3.3 Early Projects Micromachined pressure sensors have been made since 1962, when Tufte and others at Honeywell made piezoresistive sensors on silicon membranes (Tufte 1962). They have been associated with medical specialities such as cardiology since at least All the recent projects aimed at making micromachined IOP sensors have used capacitive sensors, because of the low power consumption that is possible, better sensitivity and less long term and thermal drift. Resonant designs have been avoided because of the problems of using them in a liquid but they could be considered for any fresh development since techniques are now available to allow immersed operation. The two IOP sensors below are important in the development of pressure sensors embedded in the eye but did not use micromachining Collins The device designed by Collins (Collins 1967) was an inductor-capacitor (LC) tank but had poor long-term stability, according to (Van Schuylenbergh 1996a). Cooper and Beale Cooper and Beale built an applanating guard ring pressure sensor and applied it to a cadaver eye (Cooper 1977). The sensor was a passive resonant coil/capacitor arrangement, where the inductance of the coil was changed by a moving ferrite plate that acted as the applanating surface. They obtained linear changes in resonance with IOP, demonstrating the feasibility of the wireless approach. Later, together with Constable, they reported the results of trials in vivo with dogs and rabbits (Cooper 1979). Accuracy was not found to be sufficient for use with humans, and the repeatability of readings over time was less than ideal. 3.4 Sweden - Uppsala This project is generally credited with rekindling interest in intraocular pressure sensors during the last decade thesis_djh_gdd.doc

33 Pressure Sensors Who The work was carried out in the electronics department of Uppsala University in Sweden and the Institute of Ophthalmology at the Academic Hospital there. It was funded by the Swedish Board for Technical Development (NUTEK) and first reported in (Bäcklund 1990). What Rosengren, Bäcklund and colleagues describe a pressure sensor, intended for implantation in the eye, that is very simple in concept as sketched in Figure 6 (Rosengren 1992). It consists of a microfabricated pressure-sensitive capacitor and a small handwound coil connected in a resonant tank circuit. An external device is used to detect the resonant frequency of the circuit, which changes as the pressure alters the value of the capacitor. Figure 6 Passive resonant pressure sensor from (Rosengren 1992) The device was made from two silicon wafers. The top wafer is anisotropically etched from both the top and bottom to form the thin membrane, whilst the second wafer is left untouched. The two are bonded together at 1000 C and aluminium is then deposited on the top and bottom surfaces of the resulting structure, in order to form the electrodes of the capacitor. Results and analysis The design is very simple and looks dated now, with large and to some extent uncontrolled stray capacitances and resistances due to the presence of the wafers leading to a low Q (5.5) and consequent difficulty in resolving the resonant frequency. There was also a problem with long term stability, illustrated by the curve in Figure 7. Nonetheless, this project was important in the development of pressure sensors and in generating interest in implanted sensors in the eye. thesis_djh_gdd.doc

34 Pressure Sensors Figure 7 Difference between the sensor signal and a reference barometer over a period of 135 days from (Rosengren 1992) 3.5 Germany Aachen, Köln, Duisberg Collaboration between several German institutions has developed an implantable pressure sensor that will shortly start clinical trials. Who Four institutions two medical and two engineering were involved in the project: - Department of Ophthalmology, University of Cologne, - Fraunhofer Institute of Microelectronic Circuits and Systems, Duisberg (IMS), - Institute for Materials in Electrical Engineering I, RWTH Aachen (IWE-1), - Institute of Pathology, RWTH Aachen and so were three companies: - Acritec GmbH, a manufacturer of intraocular lenses - Bytec GmbH, a software developer, and - mesotec GmbH, a new company that owns rights to this sensor. Professor Mokwa of IWE-1 at Aachen also runs the Fraunhofer IMS. He and Schnakenberg, also of IWE-1, published a survey of the considerable number of medical microsystems developed there since 1990 (Mokwa 1998), including blood pressure and flowrate system, blood hematocrit (red cell concentration) measurement, an IOP sensor and a retinal sensor The developments have been based on CMOS electronics produced by the IMS and a surface-machined capacitive pressure sensor. IWE-1 fabricates the microcoils, bonds the coils to the sensor, encapsulates in silicone and coats with Parylene thesis_djh_gdd.doc

35 Pressure Sensors The Institute of Pathology tests the cytocompatibility and long-term stability of materials and components, whilst the Department of Ophthalmology provides expertise for trials and is responsible for dimensioning the device. The project OPHTHAL was supported by the state of North Rhine-Westphalia and the subsequent IODS by the German ministry of Education, Science, Research and Technology (BMBF). Funding was also provided by European ESPRIT programmes. Publications (Schnakenberg 2000) describes the sensor and provides other references with further detail of the pressure sensor and results. (Stangel 2001) gives more detail of the implant s construction and circuitry and (Ullerich 2000) describes the micro coils. Initial animal trials with the early prototypes are reported in (Walter 2000a); the results of later trials have not yet been published but have been submitted to Ophthalmology. What The overall concept of the IOP sensor system is illustrated in Figure 8. Figure 8 Overall concept of OPHTHAL system (Schnakenberg 2000) The device itself is embedded in an artificial lens, such as is implanted during cataract operations. Labelled transponder in Figure 9, it includes three main sections. The thesis_djh_gdd.doc

36 Pressure Sensors analogue section consists of a capacitive pressure sensor, preamplifier and analogueto-digital converter. A reference pressure sensor is also included. The computational section includes a microcontroller and memory. The power and communication section uses RF induction through a coil placed around the lens. A transmitter coil is built inside a spectacle frame and linked by wire to a handheld control unit. Figure 9 Block diagram of OPHTHAL components (Schnakenberg 2000) Results and analysis This section summarises the results of the whole device. Further details of the power supply are given in chapter 7. A first prototype was made using an existing pressure sensor design, which included the analogue electronics section. This was wired to a test circuit, encapsulated in PDMS and upon testing found to have substantially the same sensitivity as when not encapsulated. Pressure readings were offset by approximately 0.01% from the unencapsulated values, and the readings were essentially linear. A second prototype included the control logic and wireless transmission mechanism, though these were not microsystems. It was encapsulated in a PDMS disc, 15 mm in diameter and 4.5 mm thick and did not leave the centre clear. The system was tested in a pressure chamber and in rabbits' eyes under general anaesthesia. It measured the intraocular pressure with at least the same precision as an applanation tonometer. The third prototype is shown in Figure 10. It incorporates a foldable micro coil and is built to sizes suitable for implantation in the human eye. This version has been thesis_djh_gdd.doc

37 Pressure Sensors implanted in vivo in rabbits for nearly a year and good pressure results have been obtained inside the PDMS encapsulation. There are some problems with the chip, however. A redesign is under way and is expected to be completed in September, when the tests will be repeated (Schnakenberg 2002). A B C Figure 10 The foldable pressure sensor Project IODS (a) Before encapsulation; sensor chip is at top. (b) After encapsulation; folded between fingers. (c) After encapsulation; squeezed by tweezers. (a) and (c) are from (Ullerich 2000). (b) is from an IWE poster. The project is continuing with the intention of making a product for clinical use. This will be produced by AcriTec, an established manufacturer of intraocular lenses, ophthalmological instruments and other products. 3.6 Germany Bremen, Hamburg, Zürich et al A collaboration between several German institutions and ETH Zürich is developing a pressure sensor that forms part of an intraocular lens to be implanted during cataract surgery. The aim is similar to the previous project but the technology is different. The collaborators include the Universities of Bremen, Hamburg and Homburg, together with ETH Zürich. Publications include (Eggers 2000a), (Eggers 2000b), (Marschner 2002), (Rehfuß 2002). Some funding is from the German ministry for education and research. Acritec Gesellschaft für ophthalmologische Produkte mbh thesis_djh_gdd.doc

38 Pressure Sensors What The sensor is based on an earlier project the Implantable Telemetric Endo-System (ITES) which is a generic implantable sensor for biomedical applications developed by 4 academic and 6 industrial partners. It is described in (Eggers 2000a). The device uses several simple chips together with a packaging technology to integrate them. The chips are a capacitive absolute pressure sensor, an oscillator-based PWM detector, and a telemetry chip that implements an inductively coupled power supply and absorption modulated telemetry. For the intraocular sensor described in (Eggers 2000b) and illustrated in Figure 11, these components are presently mounted on a 100 µm thick FR4 printed circuit substrate using flip-chip bonding. This will be replaced by a flexible 50 µm polyimide carrier. The whole module measures 6.5 mm x 9 mm and is mounted on a PMMA intraocular lens after coating with waterproofing and biocompatibility films. Figure 11 Prototype intraocular pressure sensor (Eggers 2000b) Results and analysis The device is reported as achieving an accuracy of ±2 mmhg, though the target is 1 mmhg. Prototype hand wound coils are still used and the technology involves more discrete components. The project is clearly at an earlier stage of development than the Aachen project, although it is intended to continue to clinical trials thesis_djh_gdd.doc

39 Pressure Sensors 3.7 Germany Karlsruhe There is a third German effort that is worth mentioning briefly, although it is different in nature to all the other projects. Its aim is to develop an external non-contact tonometer that can be used wherever required. The project is set at the University of Karlsruhe in co-operation with KreCo GmbH in München and is reported in (Gundlach 2000). What A novel tonometer is being developed that uses ultrasonics to excite the corneal surface and a laser interferometer to measure the extent of the induced vibration. The resonant frequency varies with the pressure behind the cornea and the IOP can be calculated from the measured frequency. Results and analysis Designs and simulations have been made and a lab prototype has been tested. Figure 12 shows how the resonant frequency varies with IOP and indicates the reproducibility of measurements, which have a standard deviation of 3.3 mmhg for the prototype. Figure 12 Resonance frequency of a pig eye from (Gundlach 2000) thesis_djh_gdd.doc

40 Pressure Sensors The goal is to produce a non-contact, low force tonometer that can be used by the layman at home. This would enable collection of better series of measurements on individual eyes without the need to implant a sensor. It will greatly assist in the diagnosis and treatment of raised IOP, if it can be made cheaply. It will be argued in section 5.1 that a separate pressure sensor can be used to control a GDD containing an active, pressure-controlled valve. It is possible that this device could be used in place of an implanted pressure sensor to control such a valve, at least for early trials, thus opening additional clinical possibilities. 3.8 Belgium KU Leuven Puers has led efforts at the Catholic University of Leuven since 1991 including two IOP sensor designs and many other capacitive sensors. He wrote a classic paper on capacitive sensors (Puers 1993). What The first IOP sensor was discontinued because full implantation was considered too invasive. The second, described here, was novel in at least two regards: the intended position within the eye, see Figure 13, and the design of the resonant circuit. Figure 13 Position of Leuven sensor The sensor is the black rectangle, arrowed, implanted under the sclera and on the surface of the choroid (Van Schuylenbergh 1996a) (Van Schuylenbergh 1996a) states that this later design is to be implanted in the outer tissue layers of the eye rather than within the eyeball itself. This approach is less thesis_djh_gdd.doc

41 Pressure Sensors invasive and consequently less risky than those designs intended for implantation within the anterior chamber or intraocular lens for example. Fabrication of the capacitive sensor also patterns the electrodes as inductors. This means that the capacitor itself forms a resonant circuit with no external components, leading to an extremely simple passive design. They note that although the inductance increases as the electrodes come closer together, it is the increase in capacitance that dominates the performance of the device. The proposed device is 3 x 5 x 1 mm and comprises a silicon membrane mounted on a glass carrier so as to enclose a thin volume. Electrodes are deposited on the inner silicon and glass surfaces to form a capacitive pressure sensor. The electrodes are patterned as coils so their inductance and capacitance form an LC tank. Sensitivity is increased by including a diode in the membrane, which produces harmonics of the resonance frequency. The harmonics are detected by another coil, allowing separation from the fundamental driving field as illustrated in Figure 14. Figure 14 Passive harmonic telemetry (Van Schuylenbergh 1996a) (Puers 2000) describes a later version built on two silicon wafers and illustrated in Figure 15, describing especially the design of the copper capacitor/inductor electrodes used in this device and gives detailed fabrication information about the electrodeposition process that has been used to make them. Results and analysis The work was verified theoretically and by a discrete prototype. Some work has been performed on the production of the integrated device and more is to be carried out. The design represents a return to the simple approach. If it proves to be successful, and there is no indication that it will not, it will produce an integrated device that is thesis_djh_gdd.doc

42 Pressure Sensors cheap to produce. It may also be easier to implant than other designs since it can be placed without opening the anterior chamber. The project is presently awaiting suitable resources to continue the work (Puers, personal communication). Figure 15 Passive resonant distributed LC pressure sensor (Puers 2000) 3.9 France Tronic s (Renard 2000) describes an industrial design for a general in vivo pressure sensor. The overall arrangement is shown in Figure 16. Who The project is a collaboration between Absys, Tronic s (industrialisation), CEA-LETI (silicon process), ESIEE, SLE, Biotronix and Versamed with funding from the EC and the French Ministry of Research and Technology. Figure 16 General arrangement of Tronic s sensor (Renard 2000) thesis_djh_gdd.doc

43 Pressure Sensors What The design is conventional using an inductive loop to power a 6.5 mm 2 ASIC containing power and data transmission circuitry and a separate sensor chip. The pressure sensor is a capacitive design, surface micromachined on wafer-bonded Silicon On Insulator material for low stress and high Q factor in deep structures. It occupies less than 1 mm 2. A reference sensor is included to compensate for temperature changes and there is also a separate temperature sensor. Results and analysis The pressure range is stated as kpa and resolution is claimed to be 20 Pa (0.15 mmhg). This is more than adequate for an IOP sensor. The power supply circuit operates at MHz and provides a regulated 3 V output. Power consumption is estimated at around 1 mw, though the paper is difficult to interpret. This design appears to be very suitable for incorporation in an IOP sensor, though it is not intended specifically for intraocular use and consequently work would be required to package it suitably Other Work For completeness, this section mentions some other projects that were not intended to produce implanted sensors but are nonetheless related. Netherlands University of Twente C. den Besten did some work on a micromachined applanation pressure sensor in the early 1990s in Elwenspoek s group at Twente. (University of Twente 1999). The work was funded by the Dutch Technology Foundation STW and the device was made but never taken to production. thesis_djh_gdd.doc

44 Pressure Sensors Sweden Umeå University There was a project at Umeå university and its hospital to make an applanation pressure sensor using a resonant design based on the change in resonant frequency a lead zirconate titanate (PZT) piezoelectric element as it is placed against the surface of the eye (Eklund 2000). US University of Minnesota Measurements by Rizq et al. confirm the feasibility of implanting a pressure sensor on the surface of the choroid, under the sclera (Rizq 2001). This is the position intended by the group at Leuven Conclusions This chapter has surveyed projects to make pressure sensors designed to be implanted in the eye in order to measure intraocular pressure. At least one of these projects has reached a stage where it is likely to become a commercially available product for clinical use in the near future. Others are also promising but are either not as advanced or have not published their results yet. It will be contended in chapter 5 that one of these sensors can and should be chosen to act as the basis for the control system for an active-valved glaucoma drainage device thesis_djh_gdd.doc

45 4 DRAINAGE IMPLANTS Devices that are surgically inserted into the eye in order to increase the outflow of the aqueous are called by a variety of names. Perhaps the most common is glaucoma drainage device (GDD). Other terms include glaucoma filtration implant (GFI) and combinations including the words implant, shunt or drain. Sources Sidoti and Baerveldt wrote a review of glaucoma drainage implants (Sidoti 1994). They summarised the clinical conditions in which drainage devices have been used as well as the design and performance of specific devices and the post-operative complications that can occur. (Stürmer 1997) surveys clinical experience with GDDs as reported by various authors. As well as summarising the results and problems occurring with several types of implant, the paper discusses some aspects of their design. A comprehensive history is provided in (Lim 1998) together with a critical review of current implants and a hint of the importance of biomaterials in the future. This review also points out that evaluation of implants is difficult because of a lack of published data and incompatibilities between that data which is published, a difficulty that was noted by the FDA when trying to establish regulatory standards (Porter 1997). These difficulties may be due in part to commercial pressures. History A horsehair was used in 1907 to keep open a drainage channel made to relieve glaucoma. Subsequent efforts used silk thread and various metallic wires; all were unsuccessful because the unregulated flow produced hypotony and because there was an inflammatory response to the materials. Tubes were similarly unsuccessful because of fibrous overgrowth and blockage of the tube leading to a cessation of drainage and a reoccurrence of hypertension. A major breakthrough was made by Molteno in 1969 when he introduced the concept of a drainage plate, outside the sclera but inside the Tenon s capsule, which was fed by a tube from the anterior chamber. Current devices are variations on this theme. thesis_djh_gdd.doc

46 Drainage Implants 4.1 Survey of Current Types This section surveys some of the current types of implant to give an overview of the designs and technologies. All the devices use a tube to feed aqueous from the anterior chamber. The tubes are all silicone with inside diameter 0.3 mm and outside diameter typically 0.64 mm unless otherwise stated. Differences between the devices lie in the nature of the arrangements for drainage and pressure control. It appears from independent tests and from the published clinical results that the valves may not always open when they should, may not always close when the pressure drops and may have varying opening set points. These results may indicate problems both in design and in manufacture and claims may be overstated for commercial reasons. Further details of the evaluations may be found in (Prata 1995), (Porter 1997) and (Lim 1998). Figure 17 shows the general arrangement of a typical drainage device and its position in the eye. The proximal end of the tube is inserted into the anterior chamber of the eye through a small incision, often made with a needle. The tube then passes back along the surface of the sclera to where the drainage plate is attached to the sclera by sutures (details differ between types). A graft taken from a donor eye is frequently used to secure the tube to the sclera and prevent the proximal end pulling out. drainage plate sutured to sclera muscle surface of sclera (inside Tenon s capsule) corneal graft covering tube to hold it in place end of drainage tube inside anterior chamber cornea, iris and pupil Figure 17 Drainage device on the surface of the eyeball (Sidoti 1994) Bleb After the device is implanted, the body reacts to this foreign object and begins to encapsulate it with fibrous tissue, forming a bleb. This bleb is extremely important in the operation of the device, because it is the principal source of resistance to the flow thesis_djh_gdd.doc

47 Drainage Implants of aqueous through the device. Aqueous flows freely out of the anterior chamber, through the tube and onto the drainage plate, from where it would spread across the surface of the sclera with minimal resistance were it not for the fibrous bleb. The bleb must grow to a suitable size to regulate the pressure acceptably and this is a major feature in the success or failure of individual cases. The bleb makes valves or flow restrictors irrelevant after a while. Animal studies have shown the pressure within the bleb to be similar to that within the anterior chamber. Molteno A Molteno plate and tube is shown at approximately life size in Figure 18. Photograph (a) shows the original design whilst (b) is a later dual plate implant designed to increase the drainage area. A The drawing (c) shows the mechanism designed to avoid hypotony. The polypropylene plate has a rim at the circumference and a V-shaped ridge of the same height (visible in photo (b)). The tissue surrounding the eye presses on both walls and provides a seal until the pressure of the draining aqueous flowing through the tube is great enough to lift it away. A tube from the anterior chamber and a drainage plate placed outside the eye and more towards the rear are features present in most subsequent devices. Figure 18 Molteno implant from (Lim 1998), scale bars are 1 cm Schocket The Schocket implant, which is not commercially available, uses an encircling band instead of a drainage plate. The silicone band runs around the eyeball just behind the muscle attachments. The two piece nature of the implant has led to a higher failure rate, particularly due to blockage of the distal end of the tube. thesis_djh_gdd.doc

48 Drainage Implants White The White pump shunt has a silicone reservoir of 0.4 ml capacity situated at the distal end of the tube. The outlet from the reservoir passes onto the drainage plate through another tube (internal diameter 0.6 mm) and both tubes have one-way, pressuresensitive valves. Pressure on the reservoir, either by normal movements of the eye or by manipulation with a finger, causes aqueous to be pumped onto the drainage plate. There have been mixed clinical results due in part to the complexity of the device, which includes two valves, and it has not been demonstrated that the pump is effective once a bleb has grown and is providing resistance. The device has fallen out of use. Krupin Krupin s design is similar to Molteno s but includes a slit valve at the downstream (distal) end of the tube, which is intended to stay closed unless the pressure within the tube is above the minimum acceptable value. The slit valve consists of a cruciform cut in the otherwise closed end of the tube. The plate is made of silicone. Baerveldt Figure 19 shows a Baerveldt implant. This also has a drainage plate but the aqueous collects underneath it, between the plate and the sclera. The plate is fastened to the sclera with absorbable sutures, to provide initial resistance until the bleb forms. A small annulus, shown at (b), limits any tendency to immediate hypotony. Figure 19 Baerveldt implant from (Lim 1998), the scale bar is 1 cm Ahmed The Ahmed is another valved implant, shown in Figure 20. A top view (b) and a side view (c) illustrate the valve mechanism, which is formed from a silicone membrane folded over to form a chamber within a clip extended from the polypropylene drainage plate. The edges of the folded membrane act as a one-way valve, with its thesis_djh_gdd.doc

49 Drainage Implants opening pressure set by the polypropylene moulding. The tapering shape of the chamber is claimed to improve flow regulation by venturi action. A B C Figure 20 Ahmed implant from manufacturer and (Lim 1998). The scale bar is 1 cm OptiMed The OptiMed OGPR-1014 implant shown in Figure 21 also has a tube and silicone drainage plate but uses a different mechanism to prevent hypotony. A block of PMMA is drilled with an array of 180 small holes, each with diameter 60 µm, mounted on the plate and connected to the tube. The block is marked with an arrow in photo (a) and illustrated in drawing (b). The diameter and length of the holes are designed with the aim that at typical aqueous flow rates, a typical IOP will develop across the holes, in accordance with Poiseuille s law. A cover over the block prevents fibrous growth blocking the holes. It is understood from the distributor that OptiMed is no longer in business. A B Figure 21 OptiMed implant from (Lim 1998), the scale bar is 1 cm 4.2 Problems Surveys of GDDs that report problems include (Sidoti 1994), (Stürmer 1997), (Lim 1998), (Eisenberg 1999) and (Ayyala 2002). The most important problems that affect the design of GDDs are immediate hypotony and overgrowth of the bleb. Other problems include tube fit and blockage and valves not closing. thesis_djh_gdd.doc

50 Drainage Implants Hypotony Loss of pressure immediately after the operation is a major problem that has to be overcome. Prior to bleb formation, there must be some other resistance in the flow path. Techniques have included implanting the tube several weeks after the drainage plate, temporarily obstructing the tube either by clamping it shut or filling it with material, securing the drainage plate more securely to prevent drainage, including pressure-sensitive valves or flow resistance channels. These problem-solving techniques have themselves all had complications: variable delay in curing raised IOP, need for second operation, unreliable closure of tube, unreliable opening and closing of valves. A GDD that can prevent hypotony would be very valuable. Bleb overgrowth Excessive growth of the bleb that is, excessive encapsulation around the drainage plate can restrict the flow so much that the device becomes ineffective. Whether or not this occurs depends upon individual physiology. Relatively recent drugs such as mitomycin C slow healing so reducing the size of the bleb and are consequently widely used. Another approach that is opening up is the use of biocompatible coatings over the drainage device, but there must then be some other method of controlling the rate of drainage of the aqueous. Again, a pressure-controlled valve would be useful. Tube fit A frequent complication is leakage around the tube where it enters the anterior chamber, which leads to hypotony. The AGFID, below, seeks to solve this problem. Tube blockage Blockage of the tube entrance, and occasionally blockage along its length, are not major problems but they do occur in a significant minority of cases. The use of porous materials is discussed below as one possible way to overcome this problem. Valves not closing Tests by Prata et al, Porter et al and others have indicated that many of the current passively-valved devices do not function correctly in vivo. The valves work in air but not in liquid, especially in aqueous, and thus serve in practice as small area flow thesis_djh_gdd.doc

51 Drainage Implants restrictions rather than as cut-off valves. This leads in some circumstances to continued loss of IOP below the minimum desirable. A reliable pressure-controlled valve would overcome this problem. 4.3 New Devices This section introduces some developments that have not yet started clinical use. AGFID The Advanced Glaucoma Filtration Implant Device (AGFID) is a novel approach to the problem. It is a single piece device with several unique features as shown in Figure 22 and in the drawings of Figure 23. Most obvious is the overall shape, designed to eliminate the leakage that is common when a tube is inserted. outside Figure 22 AGFID prototype from (AGFID 2001) inside The device is inserted through the sclera into the anterior chamber. The cross-section shape is designed to match the shape of the normal slit incision and the extension on the outside is designed to fit snugly on the sclera, making for a secure, leak-free implantation. The tapered inside end makes insertion easier. Another important feature is that the device is coated with phosphoryl choline, which has been demonstrated to reduce protein adhesion. This has enabled small flow channels to be used, which would otherwise quickly become obstructed. The main flow channel 9 is large enough not to cause significant flow restriction. There is a specific flow restriction channel 10 that is designed to produce normal flow after a bleb has formed over the outside of the implant. The bleb can be much smaller than in devices with a drainage plate. thesis_djh_gdd.doc

52 Drainage Implants Taking advantage of small channels and advances in laser technique, an innovative means to protect against post-operative hypotony has been provided. The flow restrictor 10 is initially blocked by a plug 13 and an additional very small channel 15 is the only drainage immediately after the operation. This maintains the IOP whilst the bleb forms. Once the bleb has formed, the plug is ablated using a laser shone through the cornea with a gonioscope, thus enabling full drainage. The project is a collaboration between University College London (UCL), the Institute for Ophthalmology (associated with UCL), Moorfields Eye Hospital (associated with the Institute), the University of Brighton and Biocompatibles Limited. Various aspects are reported in (Lim 1998), (AGFID 1999), (AGFID 2001) and in US patent US Figure 23 AGFID drawings from US patent The work was funded by the UK Department of Trade and Industry and Department of Health Medical Implant LINK programme and the Medical Research Council. Porous materials Another approach to biocompatibility is illustrated by the device shown in Figure 24. Figure 24 Drainage device with porous anchorage from (Jacob 1999) thesis_djh_gdd.doc

53 Drainage Implants It is a conventional tube and drainage plate device, but the bottom of the drainage plate and part of the length of the tube are covered with a porous material. The intention is that there will be cellular ingrowth into this material, attaching it securely to the sclera resulting in less micromotion of the device relative to the eye and so less irritation and consequent encapsulation. Tests of this idea were performed by modifying Baerveldt implants (Jacob 1998) with expanded PTFE underneath the drainage plate and were successful in reducing the amount of type III collagen between the implants and the sclera in rabbits after 6 months. Another, different use of porous materials is as an artificial meshwork installed through the trabecular meshwork and through which aqueous can drain. The use of an expanded PTFE mesh for this purpose is described briefly in the poster (Helies 1997). The possible use of porous meshwork as an inlet filter for a drainage device will be introduced in the next chapter. Schlemm s canal implants Schlemm s canal is a small channel that surrounds the cornea, and into which the aqueous drains from the trabecular meshwork. Schlemm s canal drains in turn into the venous system, in which the pressure is 8-11 mmhg an acceptable intraocular pressure. However, openings that are made from the anterior chamber to Schlemm s canal quickly become obstructed due to healing. (Spiegel 1999) proposed a way to avoid this by implanting a simple silicone tube between the anterior chamber and the canal. Initial experiments were satisfactory and clinical trials were planned but no results have been published to date. In a recent patent, Stegmann and Grieshaber propose a specifically shaped tube to be used in conjunction with an injected medium that expands Schlemm s canal so the tube can be inserted as shown in Figure 25 (Stegmann 2002). There are other patents with similar principles but different shaped tubes. If an approach like these succeeds, it is likely to replace conventional GDDs because of its simplicity and may well eliminate the need for active valves Pa thesis_djh_gdd.doc

54 Drainage Implants Figure 25 Schlemm s canal drainage device from (Stegmann 2002) A possible weakness of both these approaches is the chance of backflow from Schlemm s canal into the anterior chamber. A check valve (one-way blocking valve) could be added if this turns out to be a flaw in an otherwise valuable technique. 4.4 Patents There have been many patents issued in this field. In general, it is difficult to review these in isolation in part because the wording of patents is frequently obscure, since it is designed to include as many possibilities as possible in case of litigation. There is also no evidence in the patent that the device is effective or can even be fabricated. In most cases, there are no other published descriptions of the designs, since this again can lead to problems in litigation. If there are published designs or published reviews of clinical trials, these have been reviewed elsewhere and the existence of a patent is incidental. If there are not, then the patents have been disregarded for review unless they are of exceptional interest thesis_djh_gdd.doc

55 5 DESIGN CONSIDERATIONS The next few chapters build on the reviews presented in the previous chapters in order to discuss options for the design of a glaucoma drainage device. This discussion covers a wide variety of topics but after reviewing the main design issues it concentrates in particular on two critical aspects: the actuator for an active valve included in the drainage device and the power supply for this actuator, which are the subject of the succeeding two chapters. This chapter starts with a survey of the main features and design options of the overall system and then proceeds to discuss specific topics in turn. 5.1 System Design Existing glaucoma drainage devices are limited in several ways. The reliance on partial wound healing the formation of a bleb as the means of flow regulation leads to a number of problems including poor control over the flow rate, postoperative complications and large explant size because of the need for a large bleb. The basic technology that is used in the construction of these devices silicone tubes and relatively simple plastic mouldings limits the design options. Biomaterials and more advanced manufacturing techniques are already being used in new devices such as the AGFID and in designs such as Stegmann s implant in Schlemm s canal. If in addition the flow rate can be accurately controlled, without depending on the resistance of a bleb, then further benefits will accrue. Inflammation and tissue growth can be reduced because there is no need for a bleb; this reduction can be achieved by biocompatible design of the implant as well as by the use of drugs. The drainage plate can be reduced in size or eliminated altogether since it no longer needs to provide the basis for a bleb. In addition, new positions for the exit from the drain can be considered, such as in the suprachoroidal space. This space can absorb aqueous without problems but cannot be considered at present since no scarring takes place there and unregulated flows would continue unrestricted. Accurate control of flow rate is thus seen to be central to a number of possible improvements to glaucoma drainage devices. This requirement can best be met by the addition of an active flow control element, controlled by a pressure sensor. The active thesis_djh_gdd.doc

56 Design Considerations control element could in principle be either a valve or a pump, but a valve has been chosen for reasons of simplicity. By contrast, micromachined pumps are complicated and still experimental those described in the literature are too large, support too little backpressure, or are designed for flow rates that are too high with leakage at flow rates similar to the total flow rate in this application. It is also doubtful that there are any clinical situations where a pump would help but a valve would not, based on previous clinical experience with the White pump. The system design is shown in Figure 26. It consists of an implant placed in the eye, a small transmitter placed near the eye and a larger external unit placed somewhere convenient about the person. The transmitter could be mounted in spectacles or handheld while the external unit could be carried on a waistband or in a pocket. External Unit Batteries Head-mounted or Hand-held Power supply transmitter Eye Implant Glaucoma Drainage Device Micro Valve Control logic & telemetry Actuator Data storage Pressure sensor Atmospheric pressure sensor Power supply receiver Control logic & telemetry Figure 26 System design for active pressure-controlled GDD The implant itself is based around several existing components. The power supply and control logic can be adapted from an existing medical implant power supply design, as will be discussed in chapter 7 and one of the pressure sensor designs previously considered can be chosen. Similarly, an existing glaucoma drainage device can provide the basic mechanical components, though it will require adaptation to accept the microvalve. The main design options for each component are considered below thesis_djh_gdd.doc

57 Design Considerations Is the pressure sensor integrated? The pressure sensor could be fully integrated with the drain or it could be a separate device. There are potential benefits to integrating the pressure sensor in terms of cost reductions and reliability improvements but for the present these are outweighed by the advantages of a separate pressure sensor, which are: - Substantial effort has been expended on pressure sensors to measure IOP, which are further developed in consequence, and it is hoped they will soon be available for clinical use on a normal commercial basis. It makes sense to use these relatively proven devices rather than develop another. - Both the pressure sensor and the drain need external power supplies and communications. Also, the pressure of clinical value is the IOP relative to the external ambient pressure and this needs an external sensor, so the pressure measurement system cannot be self contained. Certainly whilst the drain undergoes clinical trials and possibly even thereafter, it will be desirable to be able to change the IOP setpoint externally depending on individual patients circumstances. Therefore, since both subsystems need external communications, communication between them can be conducted via the external subsystem. - If initial trials use separate pressure sensor and drains, it will simplify fault finding or replacements if required. After confidence has been gained in the complete system, attention can be given to integrating the two subsystems for later higher volume production. - A best of breed approach can be taken separately to the pressure sensor and drain. Thus, using a separate, existing pressure sensor is likely to produce better functional product, faster and cheaper. Once the concept is proven, integration can be considered with the potential benefit in the elimination of duplicated components such as microprocessors and power supplies. Is the active valve integrated? Integration is used in a different sense in this subsection. Clearly, the valve needs to be integrated physically with the drain in order that the aqueous flows through a single channel and passes through the valve. However, a question remains as to whether the valve can or should be fabricated together with the rest of the drain. thesis_djh_gdd.doc

58 Design Considerations It is the view of this author that it is very unlikely to be able to achieve a fully integrated implant with the valve and the rest of drain fabricated in one process. Indeed all current drains are currently fabricated in multiple pieces even before the valve is considered. Consideration of the geometries involved alone make it clear why a single process is not likely to be successful and additional thought as to the range of materials needed reinforces this view. However, there are advantages to counter this disappointment; accepting this cost provides a major advantage in flexibility. It is likely that the valve design can be developed independently from the drain of which it is to form a part, at least until final variations for integration. This means that the same valve design can probably be treated as a component and used in a variety of GDDs. For example, the valve could be integrated into the drainage plate of a Molteno or similar GDD, or it could be embedded inside the body of the AGFID. Such a plan gives more freedom for experimentation to find better designs. It also allows reuse of existing design and engineering, thus reducing both time and cost to completion of a functioning device. Accordingly, this plan is recommended. 5.2 General Requirements This section gathers some specific constraints that the device must meet. Size The device has to meet overall requirements on size: it must be of a suitable size to fit where required in the eye, and it must be possible to insert through surgical incisions. Surgical incisions tend to be made as small as possible to minimise the trauma and surgically-induced post-operative astigmatism. Specific advice is that the device should be 1 mm x 1 mm x 3 mm or at worst 1 mm x 3 mm x 3 mm (McNaught, personal communication) if implanted in isolation. However, it should be noted that replacement intraocular lenses are larger than this, even when folded, so detailed clinical advice should be sought about particular proposed configurations. It would be possible for the device to be larger if implanted at the same time as cataract extraction, providing that an adequate sized optic (of the intraocular lens implant) was part of the device or separately implanted thesis_djh_gdd.doc

59 Design Considerations Biocompatibility Biocompatibility will be discussed later in this chapter. Briefly, the device must not cause adverse reactions by the body and must itself survive the corrosive saline environment in which it is immersed. The standards to be used are those for a longterm implant, not for an acute or disposable device. The device must be fabricated in such a way that at least one standard sterilisation technique can be applied. Blockage Blood cells, proteins and other particles are found in the aqueous sometimes, especially after an operation such as that to position an implant. The device should be designed to reduce the chances of blockage by proteins and other materials obstructing the drain. Cost Different factors are at work in regards to the pressure sensor and the drainage device itself, so these items are best considered separately. This possibility is another benefit of dealing with physically separate devices. The pressure sensor is a diagnostic instrument and is ultimately intended for large scale implantation, unless alternative diagnostic techniques of equal merit are found. It is likely that initial usage of the device will be in eyes that are undergoing intraocular surgery. Consequently, the sensor must not significantly increase the cost or risk of the operation used to implant it (usually small-incision phakoemulsification cataract extraction with intraocular lens implantation). Thus attention to cost reduction should be a significant item in the design of the pressure sensor. The drain by contrast is a treatment and so more valuable. The cost of a drain that functions for a patient s lifetime has to be compared with the cost of alternative medical, laser and surgical modalities utilised for the same period of time. Another cost of not treating the condition is potential blindness so a successful treatment is highly prized. One or more specific operations will also be scheduled to implant the device so the cost of the device is less significant when considering the total cost of treatment. The market is sensitive to cost comparisons against traditional devices, so additional costs must still be justified by increased benefits. In summary, cost should not be a major consideration in early versions of a device. Reduced costs should be sought once significant clinical benefit is demonstrated with a design. thesis_djh_gdd.doc

60 Design Considerations Reliability The goal must be to make the implant last as long as possible, ideally decades. A faulty drain means another operation, or eventual blindness, for the unfortunate patient. However, the failure rate of current drains is surprisingly high; the reasons were examined in section 4.2. So the implant might offer some benefit even if it only had an expected life of a few years. One specific requirement is that the valve must not fall open if it fails. That is, the failure modes should be designed such that failures lead to the valve staying at its current position where possible and moving to the closed position if that is not possible but never moving to the fully open position. This is because depressurisation of the eye (hypotony) causes more rapid physiological damage than elevated IOP (hypertony). Pressure The implant must maintain the IOP within the range 5 mmhg to 20 mmhg at all times, from as soon as it is implanted. Long-term, the device should maintain IOP within the range 10 mmhg to 15mmHg whenever practicable. The sensors must be able to measure IOP between mmhg relative to atmospheric pressure, with a resolution of at least 1 mmhg. Each sensor must be able to measure absolute pressure in the range from 75 kpa to 125 kpa and must not be damaged by exposure to pressures from vacuum to 500 kpa (aircraft pressurisation accident and maximum recreational diving depth respectively), or by an unlimited rate of change of pressure. It should be noted that the resolution of relative pressure achieved by two sensors places greater requirements on the individual sensors or on a matched pair of sensors. The pressure within the eye is not constant and varies throughout the day. It can also rise suddenly due to external pressure such as that induced by a finger rubbing the eye. If fluid is immediately drained to reduce the pressure to its normal value, then the pressure will drop below the required value when the finger is removed. The design must prevent this from occurring. Note that relatively sudden pressure rises are a 1 mmhg = 133 Pa; 5 mmhg = 667 Pa; 10 mmhg = 1333 Pa; 15 mmhg = 2000 Pa; 20 mmhg = 2667 Pa; 100 mmhg = 10.3 kpa; 75 kpa = 563 mmhg; 125 kpa = 938 mmhg; 500 kpa = 3751 mmhg thesis_djh_gdd.doc

61 Design Considerations feature of some types of glaucoma (e.g. glaucomatocyclitic crisis/posner-schlossman syndrome, pigment dispersion syndrome) and these must be counteracted. Temperature The implant must work correctly within a temperature range of 35 C 45 C and should have as wide a storage temperature range as is practicable. Measurement of the temperature within the eye is clinically useful if it is available, because it provides a continuously logged record of the body temperature that is valuable in diagnosing and treating many unrelated conditions. Drain and valve It is desirable for the drain to be able to close completely or to drain the complete flow of aqueous, but limits that are more specific are necessary for an engineering design. The maximum flow rate of the drain shall be greater than 5 µl min -1 when driven by a pressure difference not exceeding 500 Pa. When closed, the leakage rate through the valve shall be less than 0.1 µl min -1. The valve shall be able to move from fully open to fully closed or vice-versa in less than five minutes. Although the valve must maintain its current position when the power is removed, it is possible that this will not maintain a constant IOP, even with a constant rate of formation of aqueous and an unchanging environment, perhaps because the valve design has a discrete number of stable positions or for some other reason, so it is necessary to specify a limit to the change. Flexibility (ability to change parameters) Flexibility is valuable in early versions; specifically, the ability to change algorithms as more is learned about the medical condition or in order to correct flaws in the system. Full integration is desirable in later versions for greater reliability and lower cost. Even here though, the set points may vary between individuals and over time for the same individual. thesis_djh_gdd.doc

62 Design Considerations The design should provide external control of the device, either through continuous involvement of the external system in the device settings or through variable logic within the implant (downloadable parameters or firmware). Self-test The implant should be able to test its major operations repeatedly on demand. For example, a common problem with currently used valved GDDs has been failure of the valves to open. This has occurred in the device as delivered and has also arisen after the device has been surgically implanted. It is thus important to be able to test such operation when and as necessary. The device should produce frequent regular signals to indicate to the control unit that it is functioning normally ( watchdog or heartbeat signals). Regulatory The device must meet regulatory requirements in the US and the EU and should meet regulatory guidelines as far as practicable. Orientation The orientation of the device will change as the person changes the position of their head. It is desirable that the device continues to function normally in all possible orientations of the head but some compromise is acceptable. The device must function normally in an upright and a supine position and all angles in between, with the head facing forward or turned to either side. Similarly, it must operate normally for positions between prone and upright. The device must maintain existing drainage in all other positions and should function normally if practicable (i.e. the drain must continue to work but may not adjust the flow rate when the top of the head is lower than the neck). Normal operation must resume when the head is back in a normal position. with suitable failsafe provision thesis_djh_gdd.doc

63 Design Considerations Noise Any noise produced by the valve is likely to be perceived by the patient. It is unknown whether such noise will be an annoyance or a reassuring sign that all is well. Devices should be tested for acceptable noise characteristics during development. Power supply availability It is assumed that the valve requires an external power supply to be present when it is opening or closing and similarly the pressure sensor requires an external power supply when it is taking measurements. The valve must remain at its then current position in the absence of the power supply and should continue to do so for as long as practicable but for a minimum of 24 hours. It is desirable that there is some method to power the device when the patient is asleep. This is of particular importance for the pressure sensor since 24 hour recordings are of clinical importance and since the control unit could generate an alarm to rouse the patient if it was necessary to operate the valve immediately. These topics are discussed in more detail in the following chapter. Bilateral implantation Glaucoma is usually a bilateral condition and in the majority of patients a functioning device would be required in both eyes. It would be essential that the functioning of a device would not interfere in any way with the functioning of the device in the contralateral eye. 5.3 Viscosity of aqueous The viscosity of aqueous humour was determined some time ago and is not accurately known (K.S. Lim, personal communication). In composition, aqueous is over 98% water and many researchers have used the viscosity of water when calculating flow rates. However, this is a risky strategy to adopt when constructing a finite element model of a valve as part of the design process and an accurate value should be confirmed. thesis_djh_gdd.doc

64 Design Considerations The situation is further confused by results reported in (Prata 1995). No difference was found in flow through a GDD in a test rig at room temperature and at body temperature. This is surprising because the viscosity of water reduces by approximately 30% over this temperature rise. There is thus either something very unusual in the variation of the viscosity with temperature that should be investigated and recorded or else the experiments should be re-examined to resolve the contradiction. Nevertheless, many subsequent researchers have used room temperature for experiments on flow rates through GDDs and for the design of such devices. A third factor should also be taken into account. The composition of aqueous varies between individuals and especially after operations, when extra proteins find their way into the aqueous. These changes in composition are believed to change the viscosity but little is known and more data are sought. It is therefore an attractive research project to measure the viscosity of aqueous over a range of temperatures and using samples from different individuals. Unfortunately, the volume of aqueous that can be extracted from an individual eye without causing problems is very limited; Lim suggests an upper limit of 100 µl. Conventional viscometers and rheometers cannot use such a small sample and it will be necessary to obtain the use of a microrheometer in order to undertake this work. An instrument that could be used by ophthalmologists would lead to the easiest collection of data. 5.4 Other Issues The drain itself can be divided into four parts for further consideration: entry port through which aqueous flows from the eye into the drain active valve transport through which aqueous flows from the entry port to the exit port and exit port through which the aqueous discharges to be reabsorbed. Various points are presented under these headings below. Entry port Aqueous enters conventional drainage devices by way of a small silicone tube. Tube occlusion is reported as a problem in 20% of cases (Stürmer 1997), so a means of reducing this problem would be useful. Perhaps the simplest possibility is to include a thesis_djh_gdd.doc

65 Design Considerations mesh before the start of the tube to filter out debris. A 3-dimensional open mesh with varying sizes of interconnected holes can collect more particles than a flat screen before becoming blocked and could be constructed from expanded PTFE as discussed in the next section. Transport Leakage around the tube is an important source of initial hypotony. Careful attention to the design to reduce this problem is required. One approach is that previously described by the AGFID team. Exit port The conventional location for the exit port of a drainage device is at the episcleral / subconjunctival surface, but there are two other sites that may be worth examining. Drainage into Schlemm s canal could simplify the requirements by eliminating the need for flow regulation, though there are concerns about possible scarring. By contrast, drainage into the suprachoroidal space needs very reliable pressure control, because the tissue never scars so it could only be considered once the active pressure control has proven reliable. However, the lack of scarring is a major advantage if that reliability can be achieved. These options need to be investigated in collaboration with medical researchers. If the exit port is in the conventional position on the outside of the sclera, then the drainage plate offers a convenient place to mount coils for power collection as well as the active valve. The pressure sensor too can be in this location once an integrated system is developed, or it can be underneath the sclera, in both cases avoiding the need to penetrate the eye to implant the device, except for the small drainage tube. Soft lithography The next chapter assumes that the valve itself will be made by micromachining silicon or perhaps glass. This is the most likely technology but a short description of an alternative technology is included here. This technology allows the creation of valves and other structures using soft polymers, which are attractive materials for implants because their mechanical properties more nearly match those of the body and because their cost is lower than that of micromachined silicon devices. thesis_djh_gdd.doc

66 Design Considerations Soft lithography was pioneered by Whiteside s group at Harvard (Xia 1998). The term covers a number of variations on a concept that exploits the properties of elastomeric polymers such as PDMS. The central idea is that a casting of a mould is made in the polymer, by pouring the polymer s precursors over the mould and allowing it to cure. The casting can then be peeled off and used in several ways. The most direct use of the casting is as a component in a system. For example if the mould has raised ridges on its surface, these will appear as long hollows in the casting and if they are pressed against a flat glass sheet, the hollows will form tubes through which liquids or gases can be passed. This technique is used to make micro gas chromatographs and many other microfluidic devices. Other uses of the casting include as a transfer stamp for lithographic printing of a wide range of materials. The power of the technique stems from the high fidelity of the casting and the simplicity and low cost of the technique. The casting can reproduce the mould with nanometre precision. The mould can be made using any manufacturing technique. Unger and others in Quake s group at Caltech reported making valves by multilayer replica moulding of PDMS (Unger 2000). Individual layers were cast in the normal way over moulds made from SU-8 on a silicon wafer. Multiple layers were then stacked on top of one another to form a more complex structure and the layers bonded together because of a variation in the process. The PDMS was made from a two-part precursor and in even numbered layers the mixture was made rich in one component, say A, whilst in odd numbered layers the mixture was rich in the other component, B. At the interface between the layers there was therefore an excess of both A and B which reacted to form a PDMS bond. The valves were simply made by crossing two tubes in subsequent layers. One tube carried the fluid to be controlled and the other, which was sealed at one end, was pressurised to close the valve. Unger et al found that the tube closed completely when the controlled fluid had an arc-shaped cross-section as shown in Figure 27. By contrast, there were always gaps remaining at the corners of rectangular tubes. The diagram is not to scale; typical dimensions were 100 µm wide by 10 µm high for the tubes, with a 30 µm separation between the tubes (i.e. the individual layers of PDMS were 40 µm thick). By varying tube dimensions, it is possible to run a control tube over another fluid tube without affecting the latter. It is also possible to control several fluid tubes from a single control tube. Thus, it is possible to fabricate complex logic with simple layouts thesis_djh_gdd.doc

67 Design Considerations Valved tube Control tube Control tube Valve Valved tube A B Figure 27 Simple valve cast in PDMS (a) side elevation, (b) plan view 5.5 Biocompatibility and Materials Biocompatibility is an important and specialised subject. There are varying definitions of biocompatible, depending on the date and viewpoint of those making the definition; for our purposes, there are two main characteristics: the device must be well-tolerated by the body, causing no injury, and the device must be unaffected by the environment within the eye and continue to function indefinitely. Patient health The continuing health of patients who have undergone an implantation is paramount, so the device must, for example, not be toxic or thrombogenic (likely to cause clots) or cause inflammation. In the past this has meant that the device must be inert but as knowledge increases advantages are being sought by using materials and designs that encourage specific reactions with living tissue. There are strict regulations to ensure this aspect of biocompatibility, enforced by government organisations such as the Medical Device Agency (MDA) in the UK and the Food and Drug Administration (FDA) in the US. This regulatory environment means that it is highly desirable for a medical device manufacturing company with previous experience of the process to make the necessary applications. The process also strongly encourages the use of materials that have been used in previous medical practice. If a material has not been extensively used for the same application, it will be necessary to test it against the various supranational standards. Coating with biomaterials, as discussed below, offers some help. thesis_djh_gdd.doc

68 Design Considerations Ocular implants have some advantages in this situation. The device is isolated from the rest of the body, so most problems are likely to have only a local impact and are not likely to be life-threatening. In patients who otherwise would lose their sight, the risks from treatment may be no greater than from doing nothing. Approval may be easier to obtain than in some other cases, in consequence. Survival of the device A separate problem of biocompatibility is whether the device can survive inside the body. The environment is very corrosive, consisting of warm salt water with many dissolved and suspended reagents and biologically active structures. The problem has been likened to that of designing a device to survive indefinitely and without maintenance in the sea. Each material that will be exposed to human tissues and fluids must therefore be tested for its corrosion resistance as well as its harmlessness to the body. For example, polyamides like nylon absorb water and degrade because their hydrogen bonds are destroyed in vivo, so while they are good engineering materials and have medical applications for such purposes as absorbable sutures, they are unsuitable for use on the surface of an implant. Encapsulation Encapsulation is a technique that increases the range of materials that can be used. A functional device, which may incorporate materials that are not biocompatible, is enclosed in a casing of a biocompatible material. This serves to protect the host from the materials in the device and protects the device from attack by the body. This is a very useful technique; its main weakness occurs when some part of the device must contact the tissue or body fluids in order to perform its function. Sterilisation Sterilisation is necessary to protect the patient but it restricts the materials and construction techniques that can be used in the device. There are four commonly used techniques of sterilisation: Dry heat Baking at C. This is too hot for most polymers. Notable exceptions include PTFE and silicone thesis_djh_gdd.doc

69 Design Considerations Steam Gas (autoclaving) High steam pressure at C. This is also problematic for some polymers Ethylene oxide is usually used. This is a room temperature treatment but is costly because of time it takes and there can be chemical interactions. Propylene oxide is also used, as are phenolic and hypochloride solutions. Radiation Gamma irradiation. This is another room temperature treatment and it can be done through sterile packaging. For micromachined devices, radiation sterilisation is likely to be most appropriate. Biomaterials Biomaterials offer a means to deal with the risks to patients health and to gain improved performance. The subject is very specialised so it is best to collaborate with an expert on biomaterials. It is also usually most convenient to use a previously proven material for the particular application (i.e. implantation in the eye) in order to avoid the need for extensive materials testing. Some materials commonly used are listed below. PDMS PDMS (Dow Corning Sylgard 184) is approved by the FDA and has long been used for intraocular lenses. It is gas permeable, which is a benefit for use in contact lenses but a problem when used as a membrane to isolate gases. The problem can be reduced by laminating an impermeable material such as platinum within the PDMS. It is flexible. The tensile strength and tear resistance of PDMS elastomer can be improved by the incorporation of amorphous silica or some other ceramics. The material is fabricated using a two part system, which are mixed, degassed in vacuum and vulcanised for two hours at 50 C. The mixture can be moulded or spincast prior to vulcanisation. Polyimide Polyimides are widely used for insulation and support films in the electronics industry and some have proven to be non toxic and stable in a biological environment. Stieglitz et al have developed processes to exploit a polyimide (Du Pont Pyralin PI 2611) in biomedical implants (Stieglitz 2000). The material can serve as a substrate and as an thesis_djh_gdd.doc

70 Design Considerations insulator. Many conventional micromachining techniques can be used with it and metals and other substances can be deposited and patterned on it. PMMA PMMA is used for hard intraocular lenses; it was the first polymer to be used for this purpose. It is regarded as biocompatible but it suffers from infection by bacteria and the immediate coverage of surfaces by proteins, especially albumin, so there is a need to avoid coverage by fibrinogen. It can be encapsulated in silicone or treated with coatings (see Figure 28). PTFE PTFE is widely used as a bioinert material in surgery and it has found application in ocular implants. Of particular interest is expanded PTFE (eptfe). Jacob s work on the use of eptfe to encourage cellular ingrowth was discussed in section 4.3 and she has also investigated its use as an alternative to scleral grafts. Helies et al have also used it as a drainage meshwork. Both groups tested the material and found it to have good biocompatibility. Apatite (Stürmer 1997) reported the work of Pandya et al in using hydroxylapatite as plate material in a GDD. Hydroxylapatite is a mineral, similar to the bioapatite that is one of the main constituents of bone, and is widely used in osteoplastic implants. Pandya et al found less overgrowth of their plate than is normal with a silicone plate, thus raising the possibility that smaller apatite plates could fulfil the function with less irritation to the body. Coatings Although silicone is somewhat biocompatible, it is not completely so and the growth of cells along the wall of the tube (fibrosis) means that implants have to be replaced every few years. This is a strong reason to improve the biocompatibility of implants. Another reason is to control the outflow resistance of the device. It was explained earlier that the resistance to fluid flow in most GDDs is largely provided by the fibrous growth over the drainage plate of the device. By choosing appropriate thesis_djh_gdd.doc

71 Design Considerations materials and coatings, it is possible to affect this growth and thus change the resistance of the GDD. For example, a surface that led to less fibrous overgrowth might provide less resistance per unit area and so enable the use of a smaller drainage plate. (Allan 1999) describes various factors and includes the picture below showing the improvements that can be achieved by coatings, in this case phosphoryl choline from Biocompatibles Ltd. A B Figure 28 The effect of coating PMMA (a) coated with phosphoryl choline, showing consequent reduction in cell adhesion, (b) without. (Allan 1999) (Hanein 2001) describes the use of polymerised tetraglyme (pp4g) to reduce protein adhesion on structures built from silicon. The structures were first passivated with silicon nitride, which is more corrosion resistant than silicon oxide in aqueous environments. Another commonly used coating is Parylene, which is also used to reduce friction on objects that need to be inserted into the body. thesis_djh_gdd.doc

72 Design Considerations thesis_djh_gdd.doc

73 6 VALVE AND ACTUATOR DESIGN This chapter examines the design of the active valve and especially the actuator. The requirements for the valve in this application differ considerably in several respects from the requirements that are addressed in most published work on microvalves. Specifically: - The flow rates and operating pressures are much less than are typical in published work. Indeed, the maximum flow rate is less than the leakage rate achieved in some designs. - The necessary speed of operation of this device is much slower than that targeted in most designs. Many published designs are strongly constrained by the desire to increase operating frequency to a point in the range of 1 Hz to 1 khz, whereas in this application a complete cycle could take up to 10 min (approx Hz). - The power available to this valve is severely restricted. Much published work assumes an unrestricted power supply, some squander power to achieve rapid operation. This restriction limits the choice of technologies for the actuator. - This application needs a flow regulating valve with near-continuous settings, whereas much published work is on the more common two-position valve used, for example, in pumps and in lab on a chip flow control applications. One very common design target is a check valve, which is a two position valve designed to offer minimal resistance in the forward direction and a high resistance in the reverse direction. - This valve needs to be latching, in that it must not move when power is removed. Most valve designs adopt a normal position when power is removed, either normally open or normally closed, and require continuous power input to adopt the other position. Such behaviour is not acceptable in this application. Because of these differences, the most relevant previous work is that done for the same application, though there is some other interesting work. This chapter starts with a brief review of actuator technologies, then surveys the prior work and discusses future development of this work. Finally, there is a description of particular issues in other parts of the valve. thesis_djh_gdd.doc

74 Valve and Actuator Design 6.1 Actuation Principles Many different physical phenomena have been used to construct micromechanical actuators and researchers continue to try more; the field is still young and there are many possibilities. Therefore, the review here is selective rather than comprehensive and concentrates on some of the most common types and those actuator types that most nearly meet the requirements of the application. Two desirable properties of the actuator have already been discussed low power operation and zero holding power and it is useful to introduce a third at this point. The third property is that the actuator should have a large throw, by which is meant that the actuator should move a significant distance. In the drainage device, the valve needs to open and close a microfluidic passage large enough so as not to be obstructed by particles in the flowing aqueous. A clear passage of diameter 50 µm is a reasonable choice for the valve opening, leading to a requirement for the throw of the actuator to be at least this large. Piezoelectric actuators The reverse piezoelectric effect is commonly used in macro-actuators and is increasingly finding favour in microactuators. When a voltage is applied across two electrodes deposited on either side of a slab of piezoelectric material, as shown in Figure 29, the piezoelectric material changes its shape slightly. When properly engineered, actuators built from these materials can be very precise, can create large forces and consume only small amounts of power since they are essentially electrostatic devices. Unfortunately, they require high voltages to operate (typically hundreds of volts) and generate very small movements (typically fractions of a percent extension). There are also some issues with materials and fabrication, so on balance these actuators are not considered further here. + displacement Piezoelectric material - thin conductive electrodes Figure 29 Principle of piezoelectric actuator thesis_djh_gdd.doc

75 Valve and Actuator Design Heated actuators Several types of actuator use heat to achieve their motion. These include bimetallic actuators (or more generally, bimaterial), shape memory actuators and actuators that use heated fluids. These actuators suffer from heat loss as the ratio of surface area to volume increases at microscale and this heat loss has to be replaced by additional input power. The losses can be reduced by encapsulating the actuator in a vacuum but over time outgassing from the surrounding materials provides a conduction path and increases the losses again. So actuators based on heating are not considered further. Direct liquid drive actuators There is a class of actuators that generate forces directly in the body or surface of the liquid, including electrohydrodynamic, electrokinetic, electroosmotic and other mechanisms. However, they are all strongly dependent on the electrical and mechanical properties of the transported liquid, which makes them complex to use in an environment where the composition of the liquid can change over time, is not homogeneous and includes particulates. Further, the applied electric fields can cause unwanted side-effects on the liquid, such as electrolysis. These techniques are therefore not considered promising candidates in this application. Magnetic actuators Magnetic actuators, also known as electromagnetic actuators, use a current in a coil of wire to generate a magnetic field and attract a piece of magnetic material. The principle of a magnetic actuator is shown in Figure 30, which also illustrates how a latching device can be made. A coil is wrapped around a soft permeable core and positioned close to a permanent magnet. The coil and core are fixed whilst the magnet is able to move but is held away from the core by a spring. It is equally possible to hold the magnet fixed whilst allowing the core, and perhaps the coil, to move. Because the spring force increases linearly with displacement, whilst the magnetic attraction between the magnet and the core increases more rapidly, there is a pull-in distance at which the magnet will leap towards the core as the one is brought towards the other. Thus, the system has two stable states: one with the spring relaxed and the magnet at a greater distance from the core, and one with the spring extended and the magnet touching the core. No power is necessary to maintain either state. thesis_djh_gdd.doc

76 Valve and Actuator Design coil A permanent magnet core N S B pull-in distance spring Figure 30 Principle of magnetic actuator The coil enables the transition between these two states. If current is passed through the coil so as to induce a south pole at the end nearest the magnet (in the illustrated case), this will increase the magnetic force between the two and cause the magnet to move towards the core. If a current is passed through the coil in the opposite direction, so as to induce a north pole at the end nearest the magnet, the magnetic force between the two will be weakened and the spring will pull the magnet away from the core. Such a device is said to be bistable and a theoretical design for a GDD valve based on this principle is presented in (Bae 2001). However, the proposed actuator is large. The magnetic actuator is nonetheless considered unattractive in this application because of power and fabrication issues. Bae estimates that 10 ma current is required under optimum conditions, so additional energy storage devices are needed to make a working system. Assumptions made in that paper to reduce power usage by altering the setpoint of the actuator are considered unreliable. Furthermore, fabricating planar coils, magnets and even permeable cores using micromachining technologies is challenging, the more so to make an integrated design using all these components. Summary Having considered many possible types of actuator, it is time to review two that are considered more promising electrochemical bubble actuators and electrostatic actuators. It should be noted that dismissal of the other techniques does not imply that it is considered impossible that one of those techniques could be used problems can be overcome but rather that electrochemical and electrostatic actuators are viewed as presenting fewer problems in this application thesis_djh_gdd.doc

77 Valve and Actuator Design 6.2 Electrochemical Bubble Actuators Bubbles have been used as the driving force for microactuators with great success they eject the ink in many designs of inkjet printers, which is the largest commercial market for microactuators to date. The bubbles in that case are generated by heating the working liquid, causing vapour bubbles to form and expand, driving ink drops before them. Thermal bubbles have also been used in other microsystems, with heaters embedded in sealed chambers filled with working liquid. As previously described, these techniques are not suitable in low power devices because thermal losses are large at small scales; too much heat is lost by conduction through the walls of the device because of the larger surface to volume ratio at smaller scales. An alternative approach is to use electrolysis to generate the gas bubbles, in a device known as an electrochemical actuator. Principle The principle of an electrochemical actuator is simple. A voltage is applied between two electrodes in contact with an electrolyte thus catalysing reactions at the electrodes in which gas is given off. The gas is contained within a closed space and as the reaction proceeds, the pressure of the gas builds up. One wall of the enclosing cell is designed to act as a piston, which extends as the pressure increases. The best known example is hydrolysis, in which water is separated into hydrogen and oxygen gases as shown in Figure 31, forcing the right side wall of the cell to move to the right. gas oxygen hydrogen water anode + V - V cathode Figure 31 Hydrolysis cell thesis_djh_gdd.doc

78 Valve and Actuator Design Advantages Advantages of electrolysis include relatively low operating voltages, typically less than 3 V, as well as low currents. Gas can be generated as slowly as necessary or desired. The cell integrates the power and the valve opening is essentially determined by the total charge transferred over time. Therefore, in a situation like this where the rate of opening of the valve does not matter, this is a major advantage. By contrast, this does not occur for electromagnetic actuators a specific current is required for a specific time to cause pull-in. If less current is applied for longer, the actuator movement is not sufficient to reach the pull-in point and it falls back to its original position when the current stops. Electrolytic reactions can proceed at room (or body) temperature, so there are no problems with isolating hot components from the body or with heat losses, and the gas bubbles are more dense than those generated thermally so more force is available. Reactions are reversible in principle (the reaction is that of a galvanic or voltaic cell or battery), so the actuator can be caused to retract as well as extend provided there is some external force acting to restore it. The cell can be stable over quite long periods when the power is removed, which is ideal for implanted use, and the construction is simple in principle. Challenges Some practicalities intrude however and the simple cell illustrated is inadequate in several regards. Firstly, the hydrogen and oxygen are mixed together and it is energetically favourable for them to recombine slowly to form water again (that is why a voltage needs to be applied to perform the electrolysis initially). So the actuator has poor long-term position stability. Also, in this design there is no way to speed up the recombination when required to retract the actuator. The reaction products must be separated but still in contact with their respective electrodes in order to overcome both these problems. Optimising these changes implies study of the electrochemistry of various possible reactions with careful choice of electrolyte, electrode materials and designs, and most likely the use of a membrane to separate the anode and cathode. Secondly, once these complications are taken into account as well as the restrictions of micromachining technology, the fabrication is no longer as simple and careful mechanical design is required to achieve a functional and reliable cell. Finally, it is necessary to use a constant current power supply and a closed loop control system to achieve good control of the actuator position and thus the IOP thesis_djh_gdd.doc

79 Valve and Actuator Design Neagu Working at Twente with Elwenspoek and others, Neagu is largely responsible for popularising the micromachined electrochemical actuator. In several papers that were collected into her PhD thesis (Neagu 1998), she investigated the use of a micromachined electrolytic cell to reversibly generate oxygen from a copper sulphate electrolyte in such a way that it could be reabsorbed by applying a reverse potential. Her goal was identical to that here to design a pressure reduction mechanism to control glaucoma so the benefits that she sought were the same as are sought here: low power usage, no power required to hold the system at a given flow resistance, isothermal operation, and long lifetime. The resulting device is illustrated in Figure 32 and Figure 33 and the component parts are described in Neagu s words beneath the figure. Figure 32 Cross section of electrochemically actuated valve (Neagu 1998) 1. the anode: the platinum electrode should have a relatively large active area because the oxygen gas has to react at the platinum electrode to form water when the actuator is in the pressure reduction state. 2. the cathode: the copper electrode should be thick enough to ensure the required life time: during electrolysis copper precipitates on the copper electrode; when the actuator pressure is reduced, copper ions redissolve in the electrolyte solution. In this cycle some copper may be lost due to poor adhesion of the deposited metal. Copper released in this way is no longer available for the electrolysis cycle. 3. the perm-selective membrane, Nafion: should be produced by a method compatible with the micromechanical techniques, and after deposition it should maintain its properties during the subsequent steps. thesis_djh_gdd.doc

80 Valve and Actuator Design 4. the electrolyte: electrolysis should yield only one type of gas (oxygen). The solution should be compatible with other materials in the actuator. 5. the deflecting membrane: this is used to convert the gas pressure into mechanical movement. It should have a small area, not bigger than 2 x 2 mm 2 ; its deflection range is from 0 to at least 50 µm. It should also have a low diffusion coefficient for oxygen, to guarantee the life-time of the actuator. Since it is in contact with the eye fluid, it is preferable to use a biocompatible material. Alternatively, it may be covered with a very thin layer of a biocompatible material (glass, metal, polymers). 6. bonding: the wafers should be bonded to each other to create a gas-tight cavity. This cannot be done with conventional high temperature bonding techniques, because this would destroy the Nafion membrane. Another alternative has to be found. Polymer bonding may be a good solution, although the swelling of the polymer when brought into contact with water can be a problem. 7. sealing: the filled cavity has to be gas-tight; the electrolyte can be enclosed in the cavity during the bonding or after bonding in a vacuum chamber, after which the cavity may be sealed with a polymer or an electrochemical plug. The Nafion semi-permeable membrane is worthy of additional comment. It allows positive ions to pass but stops negative ions including oxygen and non-polar molecules such as oxygen gas. It serves to prevent the generated oxygen from reaching the cathode after the power is turned off. Without this protection oxygen would diffuse to the cathode and recombine with the copper, causing the pressure in the actuator to be lost. The polyimide structure helps to hold the Nafion in place. A B Figure 33 Construction of Neagu s actuator (a) cross-section showing Nafion membrane and polyimide film that supports it and insulates edges of Pt anode, (b) photo showing C-shaped Pt anode surrounding Cu grid cathode (Neagu 1998) thesis_djh_gdd.doc

81 Valve and Actuator Design Neagu obtained promising results but further work remains to be done. The actuator was measured to consume 7 µw for 100 sec to create 2000 Pa pressure and at an operating voltage of 1.6 V and current less than 50 µa (giving a power of 80 µw) was able to generate a pressure of 200 kpa. The actuator is therefore capable of controlling a valve. However, there are still problems with irreversible loss of oxygen by reaction at the cathode, and with gas-tight low temperature sealing, so the device is not yet practical for use in a GDD. Another difficulty with the prototype is that copper sulphate is toxic, however the electrolyte may be changed in the process of resolving the oxygen loss so it is premature to consider this risk in detail. Subsequently, (Neagu 2000) has described an experimental apparatus that was built to explore different electrode and electrolyte materials, as well as an electronic feedback controller that enables accurate closed-loop control of the pressure in the cell. Papavasiliou Papavasiliou has reported work at Berkeley on bubble-actuated planar microvalves in a number of papers that have contributed to his PhD thesis (Papavasiliou 2001). The focus was making planar valves where all fluid flow and mechanical movement take place in the plane of the device, with the goal of simplifying the integration of multiple components and reducing costs. The designs are not directly suitable for use in the present application, chiefly because the actuating fluids and the liquid under transport are not separated by a membrane but are allowed to mix (a particle-free solution is assumed). However, the work does contain some relevant points of interest. One such point is the fabrication technique used to join wafers, which was developed to overcome some shortcomings found by previous experimenters using SOI wafers and is shown in Figure 34. The new technique is called Silicon On Epoxy (SOE). Briefly, circuitry such as electrodes is fabricated on the surface of the handle wafer, which is then bonded to the thinner device wafer using a layer of spun-on epoxy. The device wafer is then etched using DRIE to produce channels and potentially free floating structures, which are freed by etching the epoxy using an oxygen plasma, etch. A cover wafer, which can be glass, is then bonded on top using a reflowed 2000 Pa = 15 mmhg thesis_djh_gdd.doc

82 Valve and Actuator Design photoresist bond and another plasma etch is used to clean photoresist from the microchannels. The chief advantage of this technique is the ability to fabricate circuitry underneath the channels, leaving the possibility of a transparent, unmachined cover. It is not known if this technique has been further developed. (h) (g) (f) (e) (d) (c) (b) (a) { (i) Figure 34 Cross-section of Papavasiliou s valve showing from bottom (a) the thick supporting handle wafer, (b) electrodes fabricated on the surface of the handle wafer, (c) epoxy planarising and bonding layer, (d) photoresist adhesive layer, (e) fixed structure and (f) floating valve gate in thin device wafer, (g) photoresist gasket attaching (h) transparent cover. Further details of electrodes are given at (i). (Papavasiliou 2001) Other experiments concerned the shape and material used for the electrode design. These experiments demonstrated experimentally that polysilicon is not a suitable electrode material because of oxide formation and higher resistivity than metals, that the bubbles nucleate on the edges of electrodes and that designs such as interdigitated electrodes that maintain a constant potential difference between adjacent anode and cathode sections are superior to those that do not (because of resistive losses) thesis_djh_gdd.doc

83 Valve and Actuator Design Papavasiliou reported problems with platinum electrodes detaching from the wafer when no adhesive layer was used because of hydrogen generation under the electrode. Chromium and titanium were subsequently used as an adhesive layer and electrolysis of these underlying layers was prevented by ensuring that all edges of the electrodes were covered by an insulator (this does not appear to have been problematic for nucleation). Neagu had earlier reported problems with the electrolysis of chromium and reported that changing to titanium solved it, but it appears from the designs of her later devices that the polyimide layer used to support the Nafion membrane also served to insulate the electrode edges, so it is not clear whether this also helped to solve the problem. Papavasiliou s actuator design, illustrated in Figure 35, used two adjacent electrodes to electrolyse water in the same space, so both oxygen and hydrogen were given off and mixed freely as described for simple hydrolysis. It was thus not possible to control the rate of the reverse reaction electrically and alternative methods to eliminate the bubbles were needed to enable the valve to return to its starting point. Figure 35 Electrochemical actuator by Papavasiliou showing moveable gate that is extended to block channel at left by bubbles generated in side compartments. In this case, the extension is prevented by another bubble previously generated to retract the gate. (Papavasiliou 2001) These experiments demonstrated the wisdom of Neagu s choice of a twocompartment cell. One method that Papavasiliou investigated was to excite combustion of the hydrogen in the oxygen by generating a spark but this failed, perhaps demonstrating that heat loss through the wall of the bubble becomes of overriding importance at small enough scales. Using heated electrodes failed to ignite combustion, instead the bubble grew in size demonstrating Charles law. Catalysis using the platinum electrodes did speed up bubble absorption but the only method of thesis_djh_gdd.doc

84 Valve and Actuator Design controlling this rate was by physically moving the bubble. These experiments thus demonstrated the fundamental importance of choosing a cell design that permits electrical control of the reaction in both directions, as stressed by Neagu, and of avoiding a design that allows two or more reaction products to mix freely. Various iterations in the mechanical design of the valve also reinforced the need for careful design with attention to friction and possible unwanted modes of movement. Other aspects There has been some other work in the field of electrochemical microactuators, which has mainly concentrated on the electrolysis of water. There has also been considerable previous work on macroscale electrolytic systems that generate and absorb gas, namely fuel cells and batteries, as well as the extensive literature on electrochemistry in general. Some of the electrolytic reactions that have been investigated for room temperature fuel cells and for industrial metal reclamation may repay careful consideration for use in the present application, among others. One particularly attractive research goal would be to identify a suitable solid electrolyte. Fabrication of a microdevice that uses a liquid electrolyte requires special steps to fill the electrolytic cell and seal it after the main fabrication. This is an expensive and unreliable process, so an electrolyte that is solid at room temperature and susceptible to some deposition method such as spin coating, sputtering or evaporation, as well as supporting a suitable gas-evolving electrochemistry, would be very valuable Other outstanding questions are how to determine the position of the actuator, how to determine whether leakage is taking place, how to extend the life of the cell and how to detect that the cell is reaching the end of its life. The position of the actuator determines the flow resistance of the valve and so in conjunction with the IOP determines the flow rate of aqueous draining through the valve. It is therefore essential to control this position accurately in order to achieve good control of the IOP. However, it is not necessary to determine the position itself, since the goal is control of the IOP it is sufficient to measure that and compare it See for example (Larminie 2000) See (Faraday 1834) or more recent standard textbooks thesis_djh_gdd.doc

85 Valve and Actuator Design against the setpoint using a conventional closed loop control system. The pressure is already measured by the IOP pressure sensor. Even so, useful design information can be obtained from direct measurements and it is worthwhile to incorporate either a pressure sensor in the electrolytic cell or a displacement sensor for the actuator piston (perhaps a capacitive sensor using conductive films incorporated in the piston membrane and the seat of the valve) or both, if these can be achieved without penalty. There are also other measurands that would be valuable to know for various purposes. The flow rate through the valve would be useful for clinical research technology has not previously admitted direct measurement of the flow rate of aqueous and there is wide variation in the estimates (Collins 1980, Kanski 1999). Micro flow sensors are available, usually based either on thermal transport effects or electrical transport. Both of these have problems the first in achieving a low power device and the second in avoiding side effects on the liquid so there is scope for research in this area. Estimates of leakage will require experiments on specific cell designs, since recombination rates will depend strongly on the geometry of the cell and the properties of separation membranes and electrodes in particular. The pressure within the cell and the net charge supplied will also need to be measured. Much work has been done in battery technology on methods for increasing the number of charge-discharge cycles that a cell can undergo before it fails and this work should be studied in attempts to improve the lifetime of the actuator. End of life can be detected by monitoring the open circuit voltage of the cell in its fully discharged condition since this decreases as the cell reaches its end of life. (Brüsewitz 2000) discusses these topics in the context of macro-scale actuators using hydrogen/nickel hydroxide chemistry and also hydrogen/silver. 16,000 charge/discharge cycles have been obtained. The paper also points out the possibility to recover energy as the gas is reabsorbed. There is also active research in the related field of artificial muscles using electroactive polymers, which may provide additional data. There are at least two other applications in which electrochemical micro-cells may find use: power storage on the device and self-assembly of 3D structures. The latter Final discharge voltage is the open circuit voltage across the cell once all the gas has been recombined Berkeley is doing work in this area in connection with microrobotics, for example thesis_djh_gdd.doc

86 Valve and Actuator Design idea is to use gas generated in an electrochemical cell as the motive force to erect outof-plane components in microsystems. Assembly of 3D microstructures is an ongoing problem and research area, which is also discussed further in the next section. Summary Electrochemical actuators appear to be the most promising technology for active valves in glaucoma drainage devices. There are other applications in which these devices could find use and a number of interesting areas of future development. 6.3 Electrostatic Actuators Electrostatic actuators are as common in microsystems as electrostatic (capacitive) sensors, because they benefit from the same favourable scaling laws, are inherently low-power devices and are relatively easy to fabricate. They are therefore attractive candidates in any application and the chief problem for use in a GDD is their need for a relatively high supply voltage. This objection can be overcome as will be discussed, so electrostatic actuators are presented here as an alternative to electrochemical ones. First, a brief introduction to these actuators is given; this is not a technical review, which can be found in standard texts such as (Madou 2002) and other references given later, but is instead a basic introduction to the concept to support the following discussion. Principle The principle of electrostatic actuation is to make use of the attraction generated between two conductors charged to different potentials (voltages). These charges need not be moving that is, there is no current flowing so the effect is called electrostatic force to distinguish it from the electromagnetic force that arises between two currents, which is the principle of magnetic actuators. The principle is illustrated in Figure 36. magnetism is fundamentally due to the current represented by circulating electrons thesis_djh_gdd.doc

87 Valve and Actuator Design + Positive charge + - Negative charge - Negative charge - F F + Positive charge F F A B Figure 36 Principle of electrostatic actuator (a) direct and (b) comb Diagram (a) shows the simplest configuration, in which two flat plate electrodes are opposed to one another and charged to different voltages, after which each experiences a force directed towards the other. Typically, one electrode will be attached to the rigid structure of the device whilst the other is moveable, perhaps suspended from flexible beams. There are problems with this design, however: The force varies inversely with the square of the separation distance, leading to the pull-in phenomenon previously described. Although the effect can be exploited, it is usually a nuisance and makes it difficult to produce proportional motion actuators. The maximum displacement is very limited, again because the force falls off as the square of the distance. Parasitic forces between the moveable electrode and other parts of the rigid structure lead to unwanted deflections. These problems are overcome using the comb actuator design shown in diagram (b). Here the substantial forces between the three moveable fingers to the right and the four fixed fingers to the left balance out, so each moveable finger is attracted equally to the fixed fingers above and below it. Instead, the actuator relies on a second-order effect, the so-called fringing fields between the ends of the fingers and the regions of the opposing electrodes beyond them. This leads to a force acting to the left on the three moveable fingers and this force stays constant as the fingers move from right to left. The force can be increased by fabricating arrays of many fingers and the displacement can be increased by fabricating longer fingers, a task that turns out to be relatively easy at microscale. So comb actuators are widely used when an electrostatic actuator is desired, although direct-acting actuators are being reexamined in some applications where the limitations can be overcome because of the larger forces they can generate (Yeh 2001). thesis_djh_gdd.doc

88 Valve and Actuator Design Fabrication Fabrication is relatively simple in principle since the only requirements are conductive materials for the electrodes (metals or doped silicon), insulators to separate them (air, silicon nitride or oxide, or polymers) and simple mechanical structures (silicon or polymers). In order to produce a concise discussion, this section focuses on the SUMMiT process developed by Sandia National Laboratories. SUMMiT SUMMiT is a sophisticated surface micromachining process that builds four polysilicon layers on top of a silicon wafer (or five layers in the SUMMiT V process). The bottom layer is used for electrical interconnections and grounding, and the other three form the mechanical structures. Each layer is planarised using CMP, resulting in the ability to design structures with much finer tolerances than would otherwise be possible. Further description of the layers and a cross-section of an example device are shown in Figure 37. A B Figure 37 SUMMiT process (a) details of layers (MMPoly are polysilicon layers, SACOX is sacrificial oxide), (b) cross-section of gear hub (Sandia) thesis_djh_gdd.doc

89 Valve and Actuator Design Significant budgets have been spent developing both the process itself and component designs that use it, both of which are available commercially, and the applications include such items as safety controls for nuclear weapons so great attention is paid to long-term continuous reliable operation. The process is fully compatible with on-chip circuitry. Sandia have done extensive work on electrostatic actuators and mechanical power transmission; they are presently developing biofluidic components and polymer-based structures as well as improving the characteristics of the existing devices. Microfluidic suitability is improved by the introduction of additional silicon nitride layers in the process (Okandan 2001). The process is thus very suitable for use in the design of a GDD valve. Out of plane actuation The Sandia devices built with the SUMMIT process are planar all the forces and movement take place in the plane of the device. This is not desirable for a valve that must isolate the actuating mechanism from the fluid being transported, since an inplane (horizontal) valve movement necessitates an out-of-plane (vertical) membrane. Constructing a vertical membrane and then sealing it both top and bottom is a challenging fabrication task and creating a design that leaves the resulting sealed membrane sufficiently flexible is more difficult still. An alternative is to design some mechanism that transfers the in-plane actuation force and displacement to an out-of-plane motion. Figure 38 shows a mechanism that erects a mirror out of the plane of the device. The mirror itself can be seen in the centre in a partially raised position, supported by another rectangular plate; the small dots are not significant and are holes used to ensure correct fabrication and operation. Small square hinges can be seen at the corners of the mirror. The right hand edge is anchored by hinges to the substrate, the two moveable plates are hinged together and the left hand edge is hinged to the actuator, which consists of a toothed rack. The rack is moved along guides by the two gears, which are turned by cranks extending from the comb actuators at the top, bottom and right of the picture. Two comb drives at 90 power each pinion. The fourth comb drive to the left is not visible, while Sandia s eagle logo is decorative rather than functional. thesis_djh_gdd.doc

90 Valve and Actuator Design Figure 38 Sandia mirror deflection system (Sandia) This design could be used to push the valve membrane upwards. Another, simpler approach to that problem may be to use a buckling beam as sketched in Figure 39. Again designed to be fabricated using the SUMMIT process, the beam is hinged to the substrate at its right end and is compressed from the left end. An arch-like clip constrains the left end to move only in the desired direction and as a result the centre of the beam is forced to buckle upwards, aided by the discontinuous butt joint at its centre (the beam is a single piece of material made from two layers of the process). The hinge can be a simple pocket as shown or could be a more sophisticated hinge as seen previously. A beam of length 400 µm could be made to buckle upward 50 µm to actuate the valve membrane by a horizontal motion at the end of less than 15 µm. actuating force clip valve membrane buckling beam hinge substrate Figure 39 Buckling beam valve actuator Multiple beams with carefully designed buckling characteristics could be used to deform the valve membrane at several points to obtain better sealing if required. The membrane and other valve components would be machined on a separate wafer and thesis_djh_gdd.doc

91 Valve and Actuator Design bonded to the actuator wafer. The valve design would use existing techniques from previous out-of-plane valve mechanisms. It must be emphasised that the essential attribute is the ability to achieve continuous positioning. There are many valves with otherwise attractive characteristics that are only bistable, such as the curling polyimide valve from Lawrence Livermore, a bistable polyimide membrane valve described in (Goll 1997) and many others with silicon diaphragms. Power supply Electrostatic actuators are very low power devices but need a high operating voltage. Present devices are typically designed to work with a drive voltage of V. This is much larger than is needed to operate microsensors or logic, or indeed an electrochemical actuator. Each of these applications typically uses around 3 V, which can be achieved with careful design of inductive loop supplies such as those designed for embedded IOP sensors. The classic way of achieving greater voltages more turns and larger coils is not applicable in this situation, but fortunately there is a power supply design approach which allows 3V to be converted to 30 V. This is the use of a voltage multiplier circuit, which is discussed in the next chapter. The necessary trade-off of current for voltage and the inevitable losses are not a problem because the current drawn by electrostatic actuators is so low. Sandia have not produced formal studies of the power consumption of their actuators but have indicated that they use low nano amps with an operating voltage around 40 V (Allen 2002). Sandia are also producing designs that work at voltages as low as 22 V as part of a concerted effort to improve the performance of their devices (Rodgers 2000). These new designs achieve greater force in a smaller area at lower voltages; one example generates approximately 12 µn in an area of 350 x 250 µm using 30 V. Force It is useful to consider the force requirement to actuate a valve membrane and the means by which electrostatic actuators might meet that requirement. Neagu achieved successful operation of her valve and recorded the information necessary to calculate the force, so this is taken as a suitable statement of the requirements (Neagu 1998). The membrane was 1.2 x 1.2 mm and a pressure of thesis_djh_gdd.doc

92 Valve and Actuator Design 40 kpa was required to deflect it 50 µm, the distance required to actuate the valve, so a force of 58 mn is required. These figures are for the corrugated silicon nitride membrane; a lower pressure was required for the polyimide membrane. Electrostatic actuators have typically been regarded as producing only low force but the recent work at Sandia has increased their output by two orders of magnitude. The modular actuator shown in Figure 40(a) is reported as producing 15 µn at 100 V and these can be arrayed as necessary. The force produced is stated to be proportional to the voltage squared, so a force of 1.5 µn can be expected at an actuation voltage of 32 V. A 4 x 7 array of the actuators can fit in an area approximately 450 x 500 µm to produce a force of 42 µn. Two such arrays at right angles can power a pinion. B A C Figure 40 Electrostatic actuator Size is 116 x 71 µm. (a) position as fabricated, (b) actuated position, (c) 3 x 3 array of actuators driving a displacement multiplier lever system (Rodgers 2000) There is still a considerable gap between the 42 µn force available and the 58 mn required, and the leverage of the out-of-plane pushrod actuation mechanism has not yet been considered. The leverage can be made 1:1 by arranging for the pushrods to be erected through 45 before they contact the valve membrane, as sketched in Figure 41(a). The sketch shows a valve fabricated from three wafers, loosely based on the arrangement of Neagu s actuator but with electrostatic actuation. The top wafer has an inlet and outlet port and connecting channel micromachined in it. The centre wafer has a 1.2 x 1.2 mm membrane above an anisotropically etched cavity; not shown is a shallower cavity to accommodate the actuation mechanism on the layer below thesis_djh_gdd.doc

93 Valve and Actuator Design A fluidic channel wafer membrane wafer rack actuator wafer outlet port pushrods membrane anchor inlet port 50 µm 500 µm 1 mm 2 x comb drives & transmission 2 x comb drives & transmission 2 x comb drives & transmission rack pushrod pushrod dotted square outlines valve membrane 2 x comb drives & transmission spare space (electronics?) 2 x comb drives & transmission 2 x comb drives & transmission 2 x comb drives & transmission 2 x comb drives & transmission B 1 mm Figure 41 Preliminary sketch of electrostatically actuated valve (a) side elevation through one actuator of three wafers with pushrods about to deform membrane, (b) plan view showing possible arrangements of components on actuator wafer. Size is 2.6 mm x 2.6 mm x 1 mm The pushrod and rack are based on the Sandia elevating mirror. Four actuators are shown in Figure 41(b), arranged around the edges of the membrane. Each is powered by two sets of comb drives, connected through gearing. Researchers at Sandia have designed step-down transmission systems to increase the force available from the actuators. Photographs of a modular design are shown in Figure 42. A step-down of :1 has been demonstrated and the transmissions are strong enough to be able to strip gear teeth. thesis_djh_gdd.doc

94 Valve and Actuator Design Figure 42 Modular geared transmission (a) 12:1 reduction unit, (b) Six of these units are cascaded together to form a single :1 transmission assembly. (Rodgers 1998) A 2-stage transmission provides a 144:1 step-down ratio likely to be adequate for this application. The two comb drives in each of the eight blocks shown in Figure 41(b) deliver power 90 out of phase so the area of only one of each pair is counted. Each is estimated to be able to deliver 44 µn so the eight pairs can deliver 8 x 44 x mn after gearing. There are losses to be considered as well but on the other hand the required force was for a silicon nitride membrane and one made from polyimide or another polymer such as Parylene with a lower modulus is likely to be more satisfactory (Stanczyk 2000). There is also further space available within a 3 x 3 mm chip area and higher gearing could be used. This approximation shows that it is feasible to design a microsystem able to generate the forces required to actuate a valve in a glaucoma drainage device using electrostatic actuators. Latching mechanisms Electrostatic actuators return to their rest positions when the power is removed, so a latching mechanism is necessary to ensure that the valve does not change position in this circumstance. There are several ways to meet the requirement thesis_djh_gdd.doc

95 Valve and Actuator Design Some actuators, such as the electrochemical ones described before, are inherently stable and retain their position with no power (though there is a long-term creep in position). Another example of this type is the worm drive, shown right, where friction together with the shallow angle of the thread of the worm prevents the spur gear from moving except when the worm is driven. However, the helical shape of the worm gear makes it difficult to fabricate at micro scales. Single gears have been made, but not as part of an integrated device. Figure 43 Worm gear (Emerson Power Transmission Corp) Another approach is to use mechanical interference and the pawl is the most common implementation of this, being both simple and easily fabricated at microscale. The principle of operation is shown in Figure 44; the rack is free to move left and right when propelled by an actuator that is not shown. It is locked in position when the actuator is not powered by a pawl formed at the end of a flexible beam. In its natural unstressed position, the pawl engages a row of teeth on one side of the rack, preventing the rack from moving. In order to allow the rack to move, a second actuator must be energised to pull the pawl away from the rack. Such mechanisms are commonly used in the devices fabricated at Sandia. anchor pawl moveable rack Figure 44 Sketch of pawl and rack Buckling beams are another way to achieve a bistable mechanism that can be used in a latch. In a simple illustration, the free end of a flexible beam is fabricated close to a rigid, immovable constraint as shown in Figure 45. It can bend up (in the diagram) freely from its rest position A but cannot bend down because it encounters the constraint. However, if it is forced past the constraint, buckling as it does so, it reaches a point B at which it remains when released. It will remain there until forced back to A. thesis_djh_gdd.doc

96 Valve and Actuator Design flexible beam A anchor B constraint Figure 45 Principle of buckling beam This principle is frequently used in micromachined devices, with many variations. Papavasiliou used buckling beams to stabilise some of his valves. Brigham Young University have a research group on compliant mechanisms that has done work in this area and Figure 46 illustrates some results of this work. A B Figure 46 Microswitch using buckling beams (Brigham Young University) Eight curved beams are used to support and constrain the commutator of a switch. In their rest position (A), they support the commutator (light coloured area shaped like an E with an extra arm) away from the two stator contacts (light coloured rectangles with projections. When forced to the right (the actuating mechanism is not clear) the beams initially must bend beyond their preshaped curve as they rotate until both ends are in vertical alignment, at which point the end-to-end length of the beam is a minimum and so the stored energy is a maximum. Beyond this point, the beams begin to extend and assist the actuator to make the commutator contact the stators, thus The two pictures do not appear to be of the same switch (compare the structures to the right), but the switches are clearly similar and are possibly adjacent identical devices in opposite states thesis_djh_gdd.doc

97 Valve and Actuator Design closing the switch (B). When actuating power is removed, the beams continue to hold the switch closed. Actuating force in the opposite direction is necessary to cause the switch to open again. The use of multiple beams serves several purposes: it locates the central beam so that it can only move left-right and not up-down, and it increases the force necessary to move the beam and also the force holding the switch closed. Interface circuitry Kensall Wise s group at Michigan have reported (Chavan 1999) the development of an integrated circuit incorporating readout circuitry for capacitive microsensors, drivers for microactuators (including electrostatic actuators) and bus interface logic. This circuit could well be useful during the prototyping of a micromachined glaucoma drainage device, though at 4.5 mm square it is rather too large for production use. 6.4 Valve Many current valve designs could be adapted for use with the chosen actuator and for incorporation within a glaucoma drainage device; the specific choice will depend on the details of the actuator and the GDD. (Koch 2000) provides a good survey as a starting point for further definition. One particular area of concern is the valve membrane and this is discussed below. Membrane For use within the body, active valves need some means to separate the actuator from the fluid being controlled. It is possible to build valves where the fluid being controlled is allowed into the actuator space, such as those of Papavasiliou described earlier. Since it simplifies the design, this is sometimes done where the fluid is a gas and even when it is a liquid, especially when the device has a short lifetime such as a medical diagnostic. However, for long-term implantation in the body this leads to several problems, including clogging of the mechanism by particles and electrical side-effects on the body fluid, such as electroosmosis and electrolysis. The need is therefore for a membrane that separates two compartments, the microchannel in which the fluid flows from the actuator mechanism, or for a seal thesis_djh_gdd.doc

98 Valve and Actuator Design where some mechanism must penetrate from one compartment to the other. It is much easier to engineer a permanent separation with a membrane, so it is highly desirable that the valve design permits the incorporation of one. PDMS is often used to fabricate membranes in microfluidic devices because it is flexible and extensible as well as easy to deposit by spin coating. However, it has relatively high gas permeability it is used in ocular applications for precisely this reason which makes it unsuitable for use as a membrane in an electrochemical actuator. Possible ways to improve the membrane material include: - laminating the membrane to reduce gas permeability (but at the risk of delamination during the lifetime of the valve), - adding clay or other materials to reduce gas permeability, and - changing the material. Alternative materials tried by Neagu include silicon nitride, which is rather brittle and stiff, and polyimide, which is widely used in the microelectronics industry and is good on temperature (for a polymer), flexibility and lack of permeability. Parylene is another alternative and (Stanczyk 2000) describes the design and fabrication of an electrochemical actuator using a Parylene membrane thesis_djh_gdd.doc

99 7 POWER SUPPLY DESIGN The power supply is a major consideration for any active implanted device and the issues in its design are considered in this chapter. The chapter starts with a summary of the design approach for the power supply, then describes the inductive loops that are used in most designs and considers some alternative technologies. Two further sections are dedicated to power supply issues: one discusses the coils in more detail and the second introduces voltage multipliers. The final section discusses communications between the implant and the external world. 7.1 Power All implants need power to operate and most need a means to transmit information to the outside world; some also need to receive information from the outside. Implants are long term so stored power cannot be used, although short-term storage is possible. The conventional design approach is to choose the power transfer mechanism first, because it poses the most significant challenges and imposes the greatest restrictions on other areas of the design. The information transmission channels are then designed to suit, often being piggybacked on the power channel. Power cannot be transmitted using wires because of the infection risk and other problems that would arise where they entered the eye, so a wireless approach is necessary. The next major problem with transferring power to an implant is the limited size of the implant and this problem is particularly acute for intraocular implants. Thus the need to optimise the receiver efficiency is a major design criterion. Two advantages can be exploited when designing intraocular implants. Firstly, the power does not need to be transmitted over a large distance; one commonly chosen option is to embed the transmitter in the frame of a pair of spectacles. Secondly, the implant makes relatively low power demands when compared to muscle-stimulating implants such as pacemakers. Inductive loops are used much more widely than any other technique for transmitting power to biomedical implants in general and intraocular implants in particular, so they are described in depth. Other techniques are summarised in a later section. thesis_djh_gdd.doc

100 Power Supply Design 7.2 Inductive loops Inductive loops are the most widely used mechanism for wireless power transmission to implants. This is because of their high efficiency, up to 80% in ideal cases, though the efficiency in an intraocular application is likely to be less than 1%. They consist of an external, primary coil and a secondary coil that forms part of the implant as illustrated in Figure 47. The two coils together constitute a loosely-coupled transformer. The primary coil is driven with an AC current from an oscillator and this induces a current in the secondary coil that is used to power the implant. The following subsections review some useful literature. This thesis restricts itself to consideration of the implant itself, so does not cover design of the external power transmission, telemetry and control unit. Class E switching-mode amplifiers are widely used and there is extensive literature. Mutual Inductance and the Coupling Coefficient The inductance of a coil, also known as self-inductance, is defined as the magnetic flux linkage per unit current in the coil. Mutual inductance, M, is defined by the flux linkage in the secondary coil per unit current in the primary coil. The maximum possible value of the mutual inductance between two coils is L L where L and L 2 are the inductances of the two coils. More usefully, M is found to be V I & where V is the voltage induced in the secondary by a current in the primary changing at rate I &. The coupling coefficient, k, between two coils is defined in terms of their mutual inductance and is a measure of the efficiency of the arrangement for energy transfer. The coupling coefficient k is given by M. L 1 L 2 Its maximum value is 1 and a well-designed macro-scale power transfer circuit might achieve 0.8. However, the small size of intraocular coils and the distance separating the secondary coil from the primary mean that coupling coefficients of 0.01 or even less are the best that can be obtained in this application thesis_djh_gdd.doc

101 Power Supply Design Donaldson and Perkins (Donaldson 1983) sets out an analysis of inductive loops, which is stated to be the first general treatment. The paper derives the desirability of series tuning for the transmitter coil and parallel tuning for the receiver coil. It considers the tuning and phase characteristics of the link and points out the degrading effect of selfcapacitance as operating frequency is lowered. It presents specific theoretical predictions and experimental confirmation for losses and link efficiency. AC Primary coil magnetic field Secondary coil Implant Figure 47 Principle of induction coils It notes the existence of a critical frequency at which both link efficiency and displacement tolerance are optimised. The coils considered are, however, some 40 mm in diameter. In (Donaldson 1986), the author supplements the previous paper by publicising the pre-existing concept of absorption modulation by which the receiver can modulate data on the power transfer link (see section 7.6). Van Schuylenbergh et al (Van Schuylenbergh 1996b) describes a technique for automatically tuning a power transfer link. The transmitter frequency has to be carefully matched to the receiver for an effective link, which can be a particular problem when the same transmitter is used with implants in multiple patients, as might occur in a hospital. Van Schuylenbergh s technique overcomes this problem by monitoring reflected power as the driving frequency is scanned. Neagu et al (Neagu 1997) derives formulae for the characteristics of a square microcoil and points out the importance of using a distributed, transmission line approach rather than a Displacement tolerance is the ability of the system to withstand misalignment of the transmitter and receiver coil positions. thesis_djh_gdd.doc

102 Power Supply Design lumped model in order to calculate values that match experiment. Coils designed with these formulae were fabricated on an oxide-passivated silicon substrate; a large element of the parasitic capacitance is due to the presence of this substrate. The coils were made of copper and two techniques were tried for their fabrication. The first method was to sputter copper over photoresist and use the lift-off technique but this produced rather thin coils. So thick photoresist was deposited over a thin chromium adhesion layer and a copper seed layer, the photoresist was patterned and an 11 µm layer of copper was electroplated. The photoresist was then removed and the adhesion and seed layers removed from the gaps between the coils by ion-beam etching. There were 112 turns, each 14 µm wide with a 20 µm gap and 11 µm high. The measured parameters matched the calculated values, providing a useful validation of the model. The coils were 4.5 mm in diameter with a resistance of 129 Ω and when driven at the resonant frequency of 3 MHz, 0.5 mw was received but with a very low induced voltage of 0.5 V. There was also only 3.5 mm separating the coil from the transmitter, a much smaller gap than is likely to be achieved in practice. Subsequently, Ullerich et al were able to improve on this performance by removing the substrate, increasing the height of the coils and enlarging the size of the coil. Ullerich et al Ullerich et al report on the coil design for the pressure sensor described in section 3.5. (Ullerich 2000) considers circular, flexible, planar coils for embedding in a soft artificial intraocular lens. The paper presents a design analysis, fabrication process, calibration measurements and some experimental results. The design analysis is concerned with optimising the number of turns, and their thickness and separation within the size constraints of the lens. It adapts the methods described by Neagu et al. The gold coils were electroplated on a polyimide film supported by a silicon substrate, using photoresist patterning. The photoresist and substrate were subsequently removed along with thin seed and adhesion films as shown in Figure 48. The coils had an external diameter of 10.3 mm and an inner diameter of 7.7 mm, allowing them to surround the optical core of an artificial intraocular lens. The coils were 20 µm thick, 60 µm wide and had a 20 µm gap between turns, of which there were 16. The number of turns is lower than optimal in terms of voltage output; this thesis_djh_gdd.doc

103 Power Supply Design was accepted in order to increase the gap, which makes the results less dependent on fabrication variations and so increases the yield. Ullerich et al stress that an important variable in designing and fabricating the coils is their self-capacitance. This should be made as low as possible to minimise dependencies of the output voltage and resonant frequency on coil thickness and separation tolerances. Encapsulation in PDMS a material commonly used for artificial lenses was successful. Figure 48 Fabrication process for flexible planar coil from (Ullerich 2001) extends the analysis (Ullerich 2000) slightly to compare the design with a smaller coil designed to be mounted directly on top of a sensor chip. This coil has an outer diameter of 6 mm and an inner diameter of 1 mm, with 50 turns each 30 µm wide, 80 µm thick and with a 20 µm gap between the turns. It has a maximum voltage output some 20% less than the preferred design. The design frequency was MHz and a receiver voltage over 20 V was achieved giving a received power of over 10 mw. Unfortunately, the transmitter coil was not separated from the receiver at all so an unrealistic coupling coefficient of 0.13 was used that is probably more than an order of magnitude greater than that which can be achieved in clinical use (i.e. k = 0.013, ~ 2 V, 0.1 mw). The sensor is stated to need 3.5 V and 280 µw, however, and the device is intended to be put into production, so presumably there are unpublished results that achieve at least this performance under realistic conditions. Eggers et al (Eggers 2000b) reports on coil design for the second German pressure sensor. This system runs at 125 khz and this lower frequency means that a larger capacitance is required. Therefore, separate SMD capacitor chips had to be employed unlike the Ullerich design where the tank capacitor is small enough to be on chip along with the sensors and logic. This design also uses slightly more power 350 µw (Eggers 2000a). thesis_djh_gdd.doc

104 Power Supply Design Permeable Cores Rehfuß et al. report on the simulation of a planar coil with a permeable layer beneath it (Rehfuß 2002), which is intended to increase both the inductance and Q-factor of the coil. The model indicates these improvements for specific combinations of the distances shown in Figure 49 with particular frequencies. The paper compares these theoretical results with an older model that gave incorrect results, but provides no experimental or other evidence to validate the new model. (Takeuchi 2002) also investigated the use of NiFe cores within a copper planar coil and concluded that these were not desirable because they reduce the Q factor. Figure 49 Planar microcoil with permeable layer from (Rehfuß 2002) Permeable cores increase the complexity of fabrication, constrain the flexibility of the coil and provide benefits that have yet to be proven in this application, so it is suggested they not be incorporated. Other points Frequency use is constrained by legislation, although there are exemptions for shortrange medical devices. In the UK, the Radio Communications Agency and the Medical Devices Agency can provide advice. (Vandevoorde 2001) compares low power and high power links and makes several useful points: Series resonant circuits are generally used for the primary coil because this works well with a voltage source, which is simpler to design and implement than the current source needed to drive a parallel resonant circuit thesis_djh_gdd.doc

105 Power Supply Design A parallel resonant circuit is universally used for the secondary coil, because the values of inductance needed for a series circuit are too large to realise in a small coil. For low power links such as the glaucoma implant, it is necessary to optimise the efficiency of the link itself to achieve a good overall efficiency. Losses in the secondary circuit can be ignored by comparison. 7.3 Coils for Power Transfer The coils are the major determinant of the power supply s performance and so repay further study. The microfabricated coils previously described have limited the performance through various characteristics. There are a number of ways in which the production of coils for power transfer might be improved and these are examined below. The position of the coil in the eye is also a concern for a GDD as compared with a pressure sensor. Coil position Several of the pressure sensor projects reviewed in chapter 3 place the coil around an artificial intraocular lens (IOL). This allows a coil diameter around 10 mm and makes the coil parallel to a transmitter coil embedded in the frame surrounding a spectacle lens. This position is not convenient for a GDD, since it would require wires passing through the capsular bag that encloses the lens. The choice of a position for the coil requires chiefly medical expertise, since it is the reaction of the tissues, ease of surgery and other clinical matters that will determine the optimum position. Several possibilities are listed here for consideration. It might be possible to position a coil in front of the lens in the anterior chamber, since some IOLs are mounted in this position using spring clips (haptics). The coil would then be close to the trabecular meshwork where the implant is likely to be positioned and its diameter could be larger than 10 mm for improved performance. Another possible position for a coil is outside the eyeball on the surface of the sclera, in a similar position to that occupied by the drainage plate of a conventional GDD. The coil would be connected to the valve by wires running along the drainage tube or body of the GDD as appropriate. This assumes a GDD design that drains to this area thesis_djh_gdd.doc

106 Power Supply Design of the eyeball rather than draining to some other area such as Schlemm s canal. This position has the same merits and problems as it does when used for a drainage plate but it also has the problem that it is not parallel to a transmitter coil around a spectacle lens so transmission efficiency would be significantly impaired unless an alternative transmitter location can be used. However, such a coil does not need an optically clear centre so more turns can be included in this area. A third possibility is to place a coil around the outside of the eyeball so that it is parallel to a spectacle lens. If it were placed in the same position as the band used in a Schocket implant the coil would have to be joined after been placed around the eye. That would require a multipole microconnector that can survive in the environment, which seems a very challenging development. Conventional coil technology The coils designed for use with pressure sensors have used silicon micromachining technology, probably because they were developed in microfabrication laboratories. However, these coils are at a crossover size that can also be fabricated with conventional technology. Small coils have long been made to fulfil needs in the electronics industry and in telecommunications in particular. It would be worth reexamining what is already available from existing suppliers as an alternative to microfabrication. It would be possible, for example, to fabricate the coils conventionally using machines to wind the coils from wire; 20 µm round wire and 20 µm by 50 µm rectangular wire are readily obtained. An important advantage of this construction is that multilayer coils can be built, increasing the inductance available. This choice might be especially suitable for production in the developing world, where there is also a large demand for cataract and other ophthalmic surgery, so an imaginative collaboration is not beyond the bounds of possibility. NMR microcoils Nuclear magnetic resonance (NMR) spectroscopy is another area where microcoils are used and it would be worth seeking benefits from applying their techniques. for example from California Fine Wire Co., Grover Beach, CA, USA thesis_djh_gdd.doc

107 Power Supply Design Figures of merit are similar to the intraocular requirement a high inductance and a low series resistance. The requirement in NMR is for even smaller coils (< 1 mm dia.) and higher frequencies (800 MHz) than for intraocular use, so the techniques are well developed. For example, researchers at the École Polytechnique Fédérale de Lausanne (EPFL) report techniques using SU-8 epoxy photoresist and electroplating that enable construction of electroplated copper coils 55 µm thick (Massin 2001). The coils can be encapsulated in the epoxy or it can be removed. These coils are being made available commercially from the spin-off company SOTEC Microsystems SA. Adoption of these LIGA-like techniques to make master formers for the electroplating would reduce costs further. Use of true X-ray LIGA would enable even thicker coils to be made if desired. Other manufacturing techniques Ullerich noted that a limit in obtaining accurate parameters in coils fabricated by electrodeposition was the variation in thickness of the deposited metal. She increased the gap between turns in order to minimise the influence of this variation, leading to a decrease in the induced voltage. It would instead be possible to use CMP to improve the consistency of coil thickness and thus permit a smaller gap between turns. An alternative fabrication technique that has been investigated by Bell Labs for the fabrication of small coils is the use of soft moulding (Rogers 1999). Rubber stamps can be used to print metal coils on curved or soft surfaces, with wire widths and gaps of 500 nm each. This technique is faster, cheaper and gives more turns per mm than the silicon micromachining discussed above, so a theoretical examination of the characteristics of the coils produced is worthwhile. 7.4 Other techniques for Power Transmission The use of inductive loops is so widespread that it is extremely difficult to find any literature that even considers alternatives. Nevertheless, this section considers some of the most likely possibilities. thesis_djh_gdd.doc

108 Power Supply Design Optical The use of light seems attractive since the cornea is transparent. Light is transmitted through the cornea in surgical procedures such as laser ablation using a device called a gonioscope. However, that is under carefully controlled and temporary circumstances and the situation is different for an implant. The power of sunlight incident around the equator is approximately 1 kw m -2 or 1 mw mm -2. So at this power level a receiver area of 1 mm 2 would be sufficient to power a device requiring 150 µw using solar cells with 15% efficiency (this is a good but achievable efficiency). The receiver would need to be larger in practice to allow for misalignment of the incoming beam and so would be close to the maximum practical in the eye. This power density is thus a reasonable estimate. Clearly it would be necessary to ensure that this light did not fall on the retina; direct sunlight will damage the retina, quite apart from being very visually intrusive. Even scattered light from a beam aimed elsewhere would be annoying. One way to avoid visual field interference is to use near infrared radiation but the cornea scatters infrared so maintaining focus would be difficult. Some devices have used infrared for communications and inductive loops for power for this reason. However, infrared does pass through skin, so it might be possible to pass a beam through the sclera or through the eyelid onto a device on the surface of the sclera, perhaps using technology like that described in (Goto 2001). Power density must be limited, however; Goto observed a temperature rise of 2.2 K with a power density of only 0.32 mw mm -2. Maintaining the beam s focus on the receiver would also provide difficulties. The obvious possibilities are to mount the transmitter either on a pair of spectacles or on a headband. Spectacles can move around and be knocked, whilst headbands can be uncomfortable so careful design would be necessary. RF Radio Frequency (RF) transmission is a misnomer in common usage. It refers to a technique of energy transmission that uses principally the electric field component of an electromagnetic wave instead of the magnetic field component used by inductive loops (E-field rather than B-field). It does not refer to transmission at a particular range of frequencies, which is what the term literally means. In fact, the inductive loops described above operate at radio frequencies, whilst any E-field transmission scheme would need to work at much higher frequencies thesis_djh_gdd.doc

109 Power Supply Design To effectively receive power using the E-field, the antenna must be of a comparable size to the wavelength, λ, of the radiation. Antennae are often λ in length, whilst 4 wires of length λ are generally considered not to radiate or receive. This provides 20 a major problem for an ocular implant where size is severely restricted, since it necessitates working at very high frequencies. If a quarter-wavelength antenna were 3 mm long, to fit in an implant, the frequency would have to be 25 GHz, a very challenging design area. There is also a potential problem with radiation heating of the tissues it passes through. Biochemical Some fuel cell researchers have speculated about the possibility of making fuel cells powered by one of the biological energy metabolism cycles. This would clearly be an ideal power source for a device implanted within a biological organism but it remains a long-term research goal and is not practical to consider for implants at present. 7.5 Voltage and Power Considerations The voltage and power obtained directly from the inductive loop are both limited. Figures obtained by various experimenters have already been given and values of 3 V and 200 µw are taken here as being typical of those that can be achieved in practice. With careful design, these values are adequate to power the necessary logic circuits and certain types of low-voltage actuators, such as the electrochemical ones discussed previously. They are inadequate, however, for some types of actuator that might otherwise be attractive for example, 30 V is a typical operating voltage for electrostatic actuators. Fortunately, the low speed of actuation required in this application can be exploited by collecting the available power over time until there is enough energy accumulated to actuate the device. It is also possible to use simple electronic circuits called voltage multipliers to increase the operating voltage, again at the expense of a reduction in operating speed. In general, the best method of storing electrical energy is in a capacitor until a point is reached at which it becomes advantageous to use alternate forms of energy storage, such as batteries. Capacitors are simple to construct in micromachined systems but the values required for energy storage can be large and therefore problematic. Electro- thesis_djh_gdd.doc

110 Power Supply Design static actuators have a fundamental advantage here, in that the actuator is itself a capacitor of precisely the right value to store the energy needed for its actuation. The other costs are some complexity and thus increased area in the circuitry and the consequent losses, which can be recovered by integrating over a longer time. (Takeuchi 2002) reported the direct powering of an electrostatic comb drive using an inductive loop but the arrangement could not be used for this application because the secondary coil was placed directly inside the transmitter coil. Voltage Multipliers Voltage multipliers are commonly used in televisions to generate the necessary high anode voltage and are used in high-voltage experimental equipment such as accelerators but they also function at lower voltages. They can be found on integrated circuits such as EEPROMs where they are used to generate the write voltage, which is greater than that used by the logic. The most common discrete circuit is called a Villard cascade and is illustrated in Figure 50. An oscillating voltage is applied to the left side of the circuit, in this case from the inductive loop, and feeds to a ladder of diodes and capacitors. C1 C3 C1 Inductive Loop D1 D2 D3 D4 L1 D1 D2 Output Output C2 C4 C2 A B Figure 50 Villard cascade voltage multiplier (a) quadrupler and (b) single stage In order to understand the circuit, consider first the single stage (b). On half-cycles when the bottom of the inductive loop is more positive, charge flows through diode D1 and charges capacitor C1 up to the peak voltage from the coil. On opposite halfcycles, current flows through C1, D2 and C2, charging C2 to twice the peak coil voltage. This is because L1 and C1 are in series and both have the peak coil voltage on them so current flows until C2 is charged to the same total voltage, less the small forward voltage drop across the diode thesis_djh_gdd.doc

111 Power Supply Design Subsequent stages, such as the second stage in diagram (a), function in exactly the same way using the oscillating voltage across D2 as the input. Consequently, capacitor C4 charges up to the same voltage as C2 and the output voltage is four times the input. Further stages can be added as necessary. The practical limit to the number of stages depends on component leakages and resistances, particularly as drawing current from the output causes ripple to appear. The time constants of the circuit also need to be such that the design current can flow through the capacitors at the operating frequency. However, in the case of an electrostatic actuator the load is itself just a capacitor and so draws a very small current once charged. Furthermore, the time constraints are very relaxed so that it does not matter if it takes several cycles for the output voltage to reach its full value. Another type of voltage multiplier is shown in Figure 51. This parallel arrangement has the advantage that smaller capacitors can be used. Its disadvantage is that each one needs to withstand the full voltage but this is not a problem at the working voltage in this application. This type of circuit is known as a charge pump. C1 C3 C5 Inductive Loop D1 D2 D3 D4 D5 D6 C2 C4 Output Figure 51 Parallel voltage multiplier (charge pump) There is a need for voltage multipliers on integrated circuits, such as the EEPROMs mentioned above. The motivation for placing such circuits on-chip is the desire to power the chip from a single, standard voltage while there is also some fundamental requirement for a higher voltage on the chip. Another example of this is levelchanging interface circuits, powered from 5 V or less logic supplies but which must drive 12 V communications circuits. The same requirement arises in microelectromechanical devices that are to be integrated with electronics, such as high frequency micromechanical resonators. An important figure of merit for these devices, the series motional resistance, can be greatly improved (reduced) by increasing the drive voltage applied to the resonator. This is a very similar situation to the electrostatic actuator supply voltage in a GDD. thesis_djh_gdd.doc

112 Power Supply Design Such charge pumps have been built for many years based on a design known as the Dickson charge pump, (Dickson 1976), which adapts the circuit shown in Figure 51 by using diode-wired MOS transistors in place of the diodes and a two-phase clock instead of the inductive AC drive, Figure 52. Figure 52 Four-stage Dickson charge pump (Wu 1998) A problem with the Dickson design has arisen as power supply voltages have steadily reduced, which is that the voltage drop across the transistors affects the efficiency of the circuit. Circuits by Wu et al (Wu 1998) and others have overcome this problem by using more complex circuits and (Shin 2000) is representative of current thinking in this direction. Shin s prototype chip is shown in Figure 53 but it uses too much power (0.23 mw) and its output voltage is limited by breakdown at 15 V. Figure 53 Charge pump by Shin unfortunately no size is given (Shin 2000) The breakdown voltage limits the maximum output voltage because of breakdown of the MOS transistors at the end of the chain. This limit has been decreasing as the It is extremely difficult to build a chain of diodes on an IC, because of coupling through the substrate thesis_djh_gdd.doc

113 Power Supply Design geometry of the transistors has been reduced by the same process improvements that have led to lower supply voltages. High voltage CMOS There are other requirements for high-voltage drive capability controlled by lowvoltage signals, so high-voltage CMOS designs do exist. One particularly relevant example is a circuit designed by Favrat et al using a standard CMOS technology and intended to power a 60 V electrostatic motor from a 1.5 V battery. This was built and tested with the motor and worked successfully. The prototype used two existing chips for the voltage multiplication and used external capacitors to avoid the poor efficiency of on-chip capacitors. For this application, the chips could certainly be integrated and it might be possible to accept the loss in efficiency in order to integrate the capacitors as well. An integrated voltage doubler stage with an efficiency of 75% is described, so a goal of transforming 3 V to 24 V might be achieved with a 4-stage cascade and an efficiency of 30%. The design issues are discussed in (Favrat 1997) and (Favrat 1998). Micromechanical switches A novel approach to solving this problem has been tried by Valizadeh in Nguyen s MEMS group at the University of Michigan, using microelectromechanical switches to replace the transistor switches as illustrated in Figure 54. These switches are electrostatically operated with a very low pull-in voltage of 1 V and a simulation of an initial design of a 20 stage charge pump built using these switches predicted an output of 34 V using a 2 V supply. Unfortunately, the first charge pump fabricated using this design did not work. The best available reference is a presentation (Valizadeh 2001) but there is also an internal report on the University s web site. It may be that achieving a stiffness great enough to avoid stiction problems is incompatible with so low an operating voltage, but many other researchers have built micromechanical switches for other purposes that could be further developed for this application. One example is work by Becher et al. at Illinois where good performance has been achieved with an actuation voltage of less than 10 V (Becher 2002). An important aspect of these switches is the emphasis placed on reliability; up to 3x10 8 operations have been achieved in tests. Careful design of the charge pump will be necessary to limit the number of operations to this level within the extended lifetime of an implant. thesis_djh_gdd.doc

114 Power Supply Design Switch length = 100 µm Gap spacing = 100 nm Control electrode width = 40 µm Switch electrode width = 10 µm Gap under the blade = 60 nm Operating voltage = 1 V Figure 54 Micromechanical switch for charge pump (Valizadeh 2001) support springs ground plane grounded switch pad (2 nd electrode under) transmission line is shorted by switch control voltage applied to 2 nd electrode Figure 55 Low-voltage RF MEMS switch The central switch pad is 150 µm by 200 µm. The S-shaped supports reduce the actuation voltage. (Becher 2002) In summary, a charge pump can be used to generate the higher voltage necessary to operate electrostatic actuators, using the low voltage available from the implant s power supply. A development of the work by Favrat et al. to optimise the design for thesis_djh_gdd.doc