The Study of Mems Acoustic Sensor for Totally Implantable Hearing-Aid System and Micropackage Technology for Implantable Devices

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THE STUDY OF MEMS ACOUSTIC SENSOR FOR TOTALLY IMPLANTABLE HEARING-AID SYSTEM AND MICROPACKAGE TECHNOLOGY FOR IMPLANTABLE DEVICES

Rui Zhang

Submitted in partial fulfillment of the requirements

For the degree of Master of Science

Thesis Advisor: Wen H. Ko

Department of Electrical Engineering and Computer Science

CASE WESTERN RESERVE UNIVERSITY

August 2011

CASE WESTERN RESERVE UNIVERSITY

SCHOOL OF GRADUATE STUDIES

We hereby approve the thesis/dissertation of

Rui Zhang ______

Master of Science candidate for the ______degree *.

Wen H.Ko (signed)______(chair of the committee)

Philip.Feng ______

Chris.Zorman ______

______

12.17.2010 (date) ______

*We also certify that written approval has been obtained for any proprietary material contained therein. Copy right Page

This thesis contains some figures and statement s from two papers published in IEEE Journals with copy right permission. They are:

© [2009 IEEE].Reprinted, with permission, from [W. H. Ko, R. Zhang, P. Huang, J. Guo, X. Ye, , D. J. Young, and C. A. Megerian, Studies of MEMS Acoustic Sensor as Implantable Microphone for Totally Implantable Hearing Aid Systems, IEEE Transaction on Biomedical Circuits and Systems 2009, and

©[2010 IEEE].Reprinted, with permission, from [R.Zhang,P.Cong,Hung-I.Kuo, W.H.Ko,” Mircopower Implantable Telemetry Device for the Study of Micro-Package Technology for Chronic Biomedical Micro-Systems”, IEEE Sensors Conference 2010, Kona, Hawaii]

Permissions from IEEE on these two papers were obtained through e mail, On July 6, 2011, from:

Jacqueline Hansson, Coordinator IEEE Intellectual Property Rights Office 445 Hoes Lane Piscataway, NJ 08855-1331 USA

Table of Contents

Preface ...... 12

Acknowledgements ...... 13

Abstract ...... 14

Part I

Chapter 1 ...... 16

Background ...... 16

Introduction ...... 17

Objective of part I ...... 19

Outline of part I ...... 19

Chapter 2 ...... 20

Theoretical modeling of MEMS displacement sensor ...... 20

Chapter 3 ...... 32

I. Sensor Design ...... 32

II sensor fabrication ...... 36

Chapter 4 ...... 40

Sensor test setup and results ...... 40

Chapter 5 ...... 48

Discussion and further sensor design ...... 48

Conclusion ...... 52

Appendix I ...... 53

Bibliography ...... 57

Part II

Chapter 1 ...... 59

Introduction and Background ...... 59

Introduction ...... 59

Background ...... 60

Objective of part II ...... 63

Outline of part II ...... 63

Chapter 2 ...... 64

Multilayer coating theory ...... 64

Theory I. The Non-hermetic Package ...... 66

Theory II: The multiple layers coating reduces defects caused failure ...... 67

Chapter 3 ...... 73

Single packaging material coating experiments ...... 73

3.1 Test set-ups for life time evaluation of packaging material in 40 and 85 oC saline solutions...... 73 3.2 Experimental procedure and results...... 76 3.3 Calibration of life time ratio between 40 oC and 85 oC accelerated tests ...... 92 3.4 Summary of the results and discussion...... 92

Chapter 4 ...... 94

Implantable telemetry unit packaged using multilayer coating techniques ...... 94

Conclusion ...... 105

Appendix ...... 106

Bibliography ...... 109

List of Tables

Part I

Table 1 summarization of mass displacement and relative displacement…………… 28

Part II

Table 1. The longest life time device, device #7, with a 381 µm MDX 4210 coated is 38 days. The #10 device with 277.5 µm MDX 4210 failed in 6 days. Its life time is much shorter than the thinner silicone coated devices, such as #6 and #9 devices. This phenomenon could be explained by the surface contamination or the defects incorporated during the coating process……………………………………………………………….79

Table 2 Test results of a2 PCB coated with multi-layer of Silicone coated in May

2009………………………………………………………………………………………79

Table 3 Test results of a1 PCB coated with Multilayer ENCAP Epoxy in 9 2008…...…82

o Table 4 Test Results of a2 PCB coated, with Multilayer ENCAP on 8/7/2009, in 40 C

Saline……………………….…………………………………………………………….83

Table 5 the test results of a2 PCB coated with multilayer ENCAP on10/15/2009, tested in

85oC ……………………………………………………………………………………..84

Table 6 Life time test results of first run experiment with Parylene coated a2 PCBs tested in 40oC saline (#: Tested under accelerated mode, converted to normal mode life time with conversion ratio 64)………………………………………………………………...87

Table 7. Summary of packaging materials and method………………………………..100

Table 8. The pulse period, width at the beginning and end of the test………………...103

List of Figures

Part I

Figure 1. Sketches of the overall arrangement of the displacement sensor attached on umbo and structure of the sensor. a) Attachment method; b) The structure of MEMS displacement sensor……………………………………………………………...……....19

Figure 2 Simplified structure sketch of the displacement sensor……………………..…21

Figure 3. a) Mass response displacement ratio to input displacement for various damping ratios; …………………………………………………………………………………………………….25

b) Phase delay of mass displacement response X for various damping ratios

……………………………………………………………………………………………25

Figure 4 a) Relative displacement ratio, y, to input displacement, A……………………27

b) Phase delay of relative displacement with various damping ratios…….…..27

Figure 5 displacement ratio and relative displacement ratio with different damping ratio……………………………………………………………………………………....29

Figure 6 Qualitatively plotted ∆C / (∆C) max, and (∆C) max is equal to amplitude of AC capacitance signal………………………………………………………………………..31

Figure.7 Umbo and PZT calibration…………………………………………..…………32

Figure 8 Structure of the displacement sensor. All dimensions are in micro-meter. …...35

Figure 9 The sensor chip is 2.2 mm by 2.2 mm square…………………………………36

Figure 10 Fabrication flow chart (A-A cross section view) ……………………………37

Figure 11 Dry etch setup illustration……………………………………………………39

Figure 12 Infrared camera photos of the device surface

a). before HF dry etching; ……………………………………………….……39

b). after HF dry etching …………………………………………….…………39

Figure 13 Sketched illustration of lab test setup of the displacement sensor……………41

Figure 14 Photos of lab test setup………………………………………………………..41

Figure 15 Block diagram of the low noise interface circuit……………………………..43

Figure 16 Voltage output of prototype displamcement sensor with 30 mg mass and comparison with the theoritical curve ( dash line )…………………………………44

Figure 17 Measured displacement x of the mass M………………………..….………45

Figure 18 Comparison of minimum detectable SPL of three …………………………47

Figure 19 a) cross-section view of the new design; b) 3-D view of the new design; c) implanted sensor setup on umbo…………………………………………………………49

Figure 20 Suggested fabrication process of the protype sensor……………………….…50

Part II

Figure.1. (a) Permeability of polymeric materials [II-6]. (b) Effectiveness of the packaging materials [II-7-II-10]…………………………………………………………65

Figure 2. Schematic of defects in a multilayer coating………………………………….68

Figure 3. Oxygen transmission rate measured at different fabrication stage of coating process on polycarbonate (PC) film, thickness for parylene-C is 2 m, and the thickness for LPCVD nitride is 30 nm [II-15]………………………………..…………………….70

Figure 4. The structure of a film on a substrate………………………………………….75

Figure 5 Normal-Mode test setup……………………………………………………….75

Figure 6. Accelerated-Mode test setup……………………………………………….....76

Figure.7 Layout of the PCB for material test experiment.

a) narrow PCB-a1;…………………………………………………...... ………77

b) wide PCB-a2…………………………………………………………………77

Figure 8. The principle of electrolysis method used to locate the leak point on the test board……………………………………………………………………………………..81

Figure 9. Lifetime test results of silicone packaged PCB; dashed line is the linear curve fitting of good device data, Red line circled data are devices failed prematurely with identified failure mode…………………………………………..……………………….85

Figure 10 Test results of ENCAP coated PCBs listed in Table 4 and 5, epresents

O O a1PCB tested in 40 C saline; represents a2 PCB, test in 85 C saline. dashed lines are the suggested trend by using Minimum Deviation and exponential curve fitting; Red line circled data are devices failed due to the wire and soldering area failure……….….…..86

Figure 11 Specialty Coating Systems (SCS) PDS 2010 Deposition System……………87

Figure 12. Parylene C film thickness is linearly proportional with the parylene dimer weight ……………………………………………………………………………………88

Figure 13. Failure mode observations of 2 µm parylene coated devices…………...……89

Figure 14. a3 type PCB test board…………………………………………………..……89

Figure 15. Accelerated test set-up for a3 type PCB coated with parylene ……………....89

Figure 16-a. Single layer Parylene test results (with 5, 7, 11 m thickness)

Survival percentage: wider gap-80%; narrower gap-67% ……………………..………..91

Figure 16-b. Multilayer Parylene test result (with 5, 5x2, 5x3, 5x4, 5x5 m thickness)

Survival percentage: wider gap-92.5%; narrowergap-80%...... 91

Figure 17. Sketch of output signal and circuit diagram…………………………….……95

Figure 18. Telemetry IC circuit layout…………………………………………….…….96

Figure 19. Sketch of assembled telemetry unit………………………………………...... 96

Figure 20. Capacitance and pulse width change…………………………………...…….97

Figure 21. Relation between pulse period and bias resistance…………….……………..98

Figure 22. Relation between pulse period and electrical potential………………………98

Figure 23. Experiment setup and block diagram of the receiver circuit………………..101

Figure 24. Photos of received low duty cycle RF signal………………………...……..102

Figure 25. The pulse duration and current consumption versus time of #10 telemetry unit…………………………………………………………………………….………..103

Figure 26. Block diagram of the telemetry circuit with RF power and remote controlling………………………………………………………………………...…….104

Preface

This thesis summarized major research works in the field of MEMS I had done in Dr.

Wen Ko’s group. There are two parts in this thesis, Part I and Part II. Part I emphasizes the research on the MEMS acoustic sensor used for totally implantable hearing aid system. Part II emphasizes the novel micro-packaging technology for micro-implantable circuits and sensors.

Some figures and statements used in this thesis are published in the two papers, which are

IEEE Transaction on Biomedical Circuit and System and IEEE Sensors 2010 seperately.

In order to avoid self-plagiarism, the authors required the reprint permission from IEEE copy right office, and the permission was granted on July 6th 2011.

Acknowledgements

My most sincere appreciation is giving to Prof. Wen Ko. His guidance to me on both academic and personal life makes this thesis possible. It is really my luck to have such an experience working with Prof. Ko, his passion to work, his love to his family and his kindness to friends, all these make Prof. Ko is a life time modal for me to follow.

Mrs. Christina Ko, she has the same deep appreciation as to Prof.. Ko from me. She looks me as her own child, and takes care of me a lot in my personal life. So even I am ten thousand miles away from my family and friends, I don’t feel lonely.

Thanks my colleagues, Dr. Hung-I Kuo, Dr. Jun Guo, Dr. Leping Bu, Dr. Peng Cong, Mr.

Ping Huang, for their inspiring discussions and technical help on finishing this thesis.

Finally, thanks my parents, 张振和 and 张晓霞 for all their supports to me.

The Study of MEMS Acoustic Sensor for Totally Implantable Hearing-aid System and Micropackage Technology for Implantable Devices

Abstract By

RUI ZHANG

PART I. A MEMS Acoustic Sensor for Totally Implantable Hearing Aid Systems

A piezoelectric vibration source was used to simulate the umbo vibration. Umbo vibration characteristics were extracted from literature and laboratory measurement data.

By directly attaching to the Umbo in the human middle ear, this displacement sensor is able to pick up the natural vibration in the ear. This displacement sensor was expected to increase the sensitivity and has a flat frequency response in the frequency bandwidth from200 to 8000 HZ. The theoretical sensor model calculation proved the feasibility of this displacement sensor used as an acoustic sensor for hearing aid. In the laboratory test, this device can detect 40-dB SPL sound in the 1–2 kHz region, with 100-Hz channel bandwidth. The results of three different acoustic sensors, including an acceleration

14 sensor are summarized and compared. A preliminary design of the implantable displacement sensor for totally implantable hearing-aid systems is also presented.

PART II. Micropackage Technology for Implantable Devices

Packaging is one of the critical processes for implantable MEMS biomedical systems.

This thesis explored a non-hermetic multiple-film-layers micro-package technology.

MDX 4210, FP 4450 and Parylene C are studied as packaging materials for the implantable micro-systems. These materials are coated on the comb figures structured

PCBs and then merged into 40OC saline for a life time evaluation or 85OC saline for accelerated life time evaluation. The failed devices are tested for failure modes. A previously designed micro power telemetry unit is packaged using this technology for evaluation. The telemetry unit is able to measure bio-potential, resistance, temperature and pressure change along with a MEMS pressure sensor. The characteristics of this telemetry unit are measured. Packaged devices were immersed in 40oC saline solution for simulated life time evaluation. The first group of devices without Parylene C coating has a life time over 70 days in 40oC saline bath.

Chapter 1

Background

Conventional hearing aids can offer moderate rehabilitation inherent limitations, such as ear canal irritations, distortion, and occasional ringing, and social problems with being perceived as handicapped, have deprived many patients of clear hearing. Partially implantable cochlear and middle ear hearing-aid systems, with external microphone and

RF electronic link, can enable those with severe hearing loss to gain improved hearing and speech function. However, the external microphone and electronics create concerns about reliability, inconvenience, and social stigma. It is, therefore, highly desirable to develop totally implantable hearing-aid systems with implantable microphones [I-1, I-2].

Several approaches of implantable microphones coupled to middle ear bones have been reported. Piezoelectric material can be used to sense the malleus vibration [I-3]. However, the material is stiff and difficult to maintain precise contact with umbo. An optic-fiber approach is complex; it consumes sizable power and also has the problem of temporarily lose signals [I-4]. These approaches all suffer from performance degradation when large shocks or sudden changes of air pressure occur. A magnetic sensor attached on malleus head was tried in Dr. Ko’s research group with encouraging results [I-5]. However, the loading effect is large, and the magnetic material is MRI incompatible. The subcutaneous microphone is being evaluated [I-6, I-7]; the noise from chewing and body movement as

16 well as the tissue growth and sensitivity degradation with time are concerns. Studies on the vibrational behavior of middle ear Ossicular bones were made. The umbo presents as the most promising site for acoustic sensor to be coupled to as the implantable microphone, because I) The sound is naturally picked up from temporal bone by the sensor without the possible attenuation or signal distortion, II) There is a 20 dB gain to the sound of the human outer ear which pre-amplified the acoustic signal for the sensor,

III) Eliminated the directivity problem of other sensing methods for implantable hearing aid system, IV) The human body act as a mechanical filter which avoided the body noise, such as the noise caused by walking or chewing food.

Introduction

A series of research work focusing on designing high performance implantable acoustic sensors coupled to umbo had been done in Dr. Wen H Ko’s group, Case Western Reserve

University, including non-contact spring coupled displacement sensor [I-8] , accelerometer[I-9] and laboratory test equipment that is correlated with the response of human temporal bones. In this thesis, a novel approach of using MEMS displacement sensor, attached on umbo, as the implantable microphone to be used in implantable hearing aid system is studied. Based on the previous research work in our lab and literatures, the requirements of the acoustic sensor are proposed as following: 1) nearly flat frequency response from 250 to 8 kHz; 2) input sound range from 40 to 100 dB SPL;

3) small size and low power consumption; 4) less than 20 mg mass and force loadings on umbo to limit the loading effect on the umbo vibration; [I-10]; 5) withstand large low- frequency displacements of umbo while sudden air pressure changes or large body shocks

17 occur. Under these situations, the umbo displacement may be up to 100–1000 nano- meters, which could be 105 times that of the umbo amplitude responding to a low input sound. Therefore, specially desiigned sensor and coupling techniques need to be developed to accommodate this large low frequency input surge. In our design, by employing the principle of traditional second order mechanical system, we developed a high pass mechanical vibration filter for a displacement sensor, with a cut off frequency of 200 Hz -- 400 Hz. Fig.1 shows the overall arrrangement of the implantable displacement sensor and a sketch of the proposed sensor structure. The objective of this thesis is to study the feasibility of the proposed high sensitivity displacement sensor used as an implantable microphone for totally implantable hearing aid systems.

b)) Figure 1. Sketches of the overall arrangement of the displacement sensor attached on

umbo and structure of the sensor. a) Attachment method; b) The structure of MEMS

displacement sensor

Objective of part I The objective of this thhesis is to sstudy the feasibility of a high sensitivity displacement sensor used as an implantable microphone for the totally implantable hearing aid system.

Outline of part I Chapter 1. Introduction and Background

Chapter 2. Theoretical model of the MEMS displacement sensor

Chapter 3. Design and fabrication of the displacement sensor

Chapter 4. Test setup and results

Chapter 5. A Proposed design of capacitive displacement sensor as the implantable

Microphone

Chapter 6. Conclusion

Chapter 2

Theoretical modeling of MEMS displacement sensor

The theoretical model of the MEMS displacement sensor is studied to verify the feasibility and the sensitivity performance over the sub-audio and audio frequency ranges.

Figure 2 Simplified structure sketch of the displacement sensor

The simplified structure of the MEMS displacement sensor is sketched as in Fig.2. There are two assumptions made in this model. First, the sensor is vertically attached on top of the umbo and aligned perfectly in the center of the umbo, so there is nearly no loss in the transition from the umbo to the sensor; and the constant gravitational force may be neglected when we only concern about the ac sound signals. Second, the umbo has an infinite driving ability to the small sensor attached, so the input excitation of the device is displacement generated from umbo. There are two masses involved in this system, one is the m, represents the diaphragm, and the other mass, M, represents the sensor substrate.

The mass M experiences three forces, force from spring, damping force and the inertia force. The inertia force F is related to the acceleration, a; if the input harmonic

* displacement is A = Asin(ωt), by applying Newton’s second law to the mass, M, we have F =Ma. And at any time t

2 dA* dx d x c( ) k(A*  x) M 2 (1) dt dt dt

Where x is the absolute displacement as measured from the mean equivalent balance position, A* is the input displacement and c is the damping factor. The spring constant is, k, and the mass is, M. The complete general solution of the displacement x (t) to the equation (1) is:

x(t) xcp (t) x (t)

x(t) C e12tt C e  Xsin( t  ) (2) 12

In equation (2), the first two terms are transient response and C1, C2, λ1 λ2 are integration constants which depend on the initial conditions. The third term is the steady state

21 response. The steady response xp(t) =Xsin(ωt+ϕ) is what we are interested, so xp(t) will be solved from equation (1) in the following steps.

dA dx d2 x c(*  ) k(A  x) M dt dt* dt 2 dx2 dx dA M c kx kA c * dt2 dt* dt dx2 dx M c kx kAsin t  cA cos t (3) dt2 dt

set kA= Nsin and cA Ncos

dx2 dx M c kx Nsin sin t  Ncos  cos t dt2 dt dx2 cdx k N xsin(t)  dt2 M dt M M fromsin 22 cos 1

N(kA)(cA)22 c where   tan1 k dx2 cdx k kA c A x()()sin(t)22   dt2 M dt M M M

kc the resonant frequency is  and damping ratio is n M 2Mk dx2 dx 2x(A)(2A)sin(t) 2222       dt2 nn dt n n dx2 dx 2x 2  22A1 (2 /  )sin(t   ) (4) dt2 nnn dt n xp (t) Xsin( t ) = Xsin  sin  t  X cos  cos  t if P=Xsin Q= Xcos then xp (t) Psin t Qcos t,substitute xp (t) into equation (4) by comparing the coefficient of sin( t) and cos( t)

22A1 (2 / ) P(kM)sin(t)nn2 (k M22 ) (c ) 2

cA1(2/)22   Qsin(t)nn (k M22 ) (c ) 2

22A1 (2 / ) x (t)  nn[(k M2 )sin t c cos t] p (k M22 ) (c )2 sin and cos can be defined as

c sin  (k M22 ) (c ) 2

(k m 2 ) cos  (k M22 ) (c ) 2

222 nnA1 (2 / ) kM  c then xp (t) [ sin t cos t] (k M22 ) (c ) 2 (k  M  22 ) (c ) 2 (k  M  22 ) (c ) 2

2A1 (2 /) 2 = n n (sin sin t cos  cos  t) (k M22 ) (c ) 2

22A1 (2 / ) = nnsin( t ) (k M22 ) (c ) 2

M ω 22A1+(2ζω / ω ) the amplitude X = nn (k - Mω 22)+(cω) 2

2 kM  n

kA 1 (2 /  )2 A  so Xn = 1 (2 )2 22 2   (k M ) (c ) (1 ( )22 ) ( 2 ) 2 n  nn

 set r= n A X1(2r) 2 (1 r22 ) (2 r) 2

X 1(2r)  2  (5) A 2 2 2 1r 2r   The unit of the term of

2 22 mA1(2/nn )kA1(2/) n is newton, so it is the equvilent exitation force applied on the mass.

The displacement ratio of mass M, to the input excitation, is plotted as Fig.3 - a.

The phase delay is denoted by  sin c 2  r tan    cos k M22 1 r where c 2 Mk 2r   tan 1 1r 2 The phase plot of X / A is shown in Fig.3- b

ζ=0.3

ζ=0.6 ζ=1

ζ=0.3

ζ=0.6 ζ=1

b) Figure 3. a) Mass respoonse displacement ratio to input displacement for various damping ratios; b) Phase delay of mass displacement response X for various damping ratios

The relative movement between mass displacement, x (t), and input displacement, A*(t), is the deformation of spring. This deformation determines the sensor capacitance output.

If this relative movement is denoted by y, then

y A* x , by substitute y into equation  1 dy d2 y dA2 ky c  m()  * dt dt22 dt dy2 dy dA2 mckym * dt22 dt dt dy2 dy 2sin() 22yA   t dt2 nn dt This equation has the same format as equation (1), so apply the same procedure as solving x, the relative displacement y can be solved.

The relative displacment amplitude is

A 2 1 Y  2  2 2 2 n 12 nn

2 Y r  2 A 2 2 12rr   (6)

The relative displacement amplitude, Y, normalized to the input excitation is the relative displacement ratio, Y/A, and it is plotted in Fig.4-a.

The phase delay, ' of y, plotted in Fig. 4-b, is derived with the following equation

3 1 2 r  'tan 22 1(4 1)r

ζ=0.3

ζ=0.6

ζ=1

ζ=0.3

ζ=0.6

ζ=1

b) Figure 4 a) Relative displacement ratio, y, to input displacement, A.

b)) Phase delay of relative displacement with various damping ratios

The displacement ratio, X/A, and relative displacement ratio, Y/A, with different damping ratio, ζ, were plotted together in Figure 5 for comparison. When the damping ratio is greater than 0.7, there is no overshot for the relative displacement, y, and the peak value of y does not occur at the resonant. It is observed that, the (X/A + Y/A) is not always equal to 1, in fact, it is greater than 1 in most of times, and X/A is always greater than 1 in frequency range below √2ωo. These results are very different than from the same system driven by constant force F sin (ωt). The analysis of the force driven system is given in the Appendix I for comparison. Table 1 summarized the features of X/A and y/A with various damping ration

Phase delay Damping Peak displacement and Peak relative displacement of X at ratio ζ frequency, X/A and frequency ,y/A resonance O 1 [email protected] ωn ~1when ω>>ωn 90

O 0.6 [email protected] ωn ~1when ω>>ωn 90

O 0.3 [email protected] ωn [email protected] ωn 90

Table 1 summarization of mass displacement and relative displacement

Damping ratio Damping ratio ζ=1 ζ=0.6

Damping ratio ζ=0.3

Figure 5 displacement ratio and relative displacement ratio with different damping ratio

This displacement sensor is really a mechanical to electrical transducer that converts the dynamic displacements to the capacitance changes. The capacitance of this displacement sensor can be written as

SS C,in this case ,C YYy0  S C0  when the input displacement, A, is zero Y0

Where ε is vacuum permittivity, Y0 is the original distance between diaphragm and substrate and S is the overlapped area of the two electrodes that forms the capacitance C.

When the substrate, M, vibrates with the input displacement, A, the capacitance also changes. ∆C can be expressed as:

SS yS CCC 0    2 YyY0000 Y Y*y yS if y << Y0 , then  C 2 Y0

The amplitude of ∆C at a given frequency is

∆C = Qy

= (∆C) max

2 Where Q = (εS/Y0 ) is a constant.

From the equation, ∆C is linearly related to relative displacement, y (t). Figure 6 plotted the predicted Co and ∆C with the damping ratio of 0.6. (Y0 = 2 µm in the design, A is of the order of 1-10 nano meter)

∆C

Figure 6 Qualitatively plotted ∆C / (∆C) max, and (∆C) max is equal to amplitude of AC capacitance signal.

Chapter 3

I. Sensor Design The frequency responsse of umbo and PZT calibration was measured by the previous lab member Ping Huang [I-8] and Mark Zurker [I-9] separately. The vibration amplitudes of the umbo and PZT stimulation unit were measured by a Laser Doppler Vibrometer

(LDV). The results are plotted in Figure 7. The 100 dB SPL will cause a maximum of

150 nm displacement on umbo at 1 KHz. At 40dB SPL, 1 KHz, the amplitude would be

0.15 nano meter. The sensor design is aimed to operate ffrom 40 dB to 100 dB SPL, at 1

KHz.

Figure.7 Umbo and PZT calibrration

The sensor, as shown in Fig. 1b, and duplicated below, is a rectangular thin silicon diaphragm supported by four cantilever beam springs on four corners anchored through

Si dioxide posts, the gap between the sensor diaphragm to the oxide thickness and is about 2 µm. The capacitor between the diaphragm and the substrate forms the capacitive sensor.

The design considerations and calculations of silicon microchip displacement sensor are outlined below. First, most of the voice frequencies in human ordinary life are in the band of 200 Hz to 8 KHz. which is the interested bandwidth of the sensor. Second, the sensor is supposed to be attached on the Umbo, without other mounting support, the mass loading effect of the umbo is a criitical factor. A proper weight of the sensor ensures the vibration can be properly converted into electrical signal without distortion or degradation. From the literature [I-8], when umbo has a mass loading greater than 20 mg, the low frequency response would be degraded. Therefore the mass of the displacement sensor (M + m) should be about 20 milligrams. If the critical frequency of the sensor, the resonant frequency n of the M and springs k discussed in Chapter 2, is set to Fc= c /2

= 200 Hz. According to equation

k   m

The spring constant k should be 31.5 N/m. Referring to Fig.1 of Chapter I, there are four folded-cantilever beams supporting the diaphragm m and M, the k value of each beam should be 7.9 N/m. For a given spring constant, the dimension of the cantilever-beams springs can be calculated from the equation below:

EBh3 k  3 4 L

Where, h, L, B are the thickness, length, and width of the spring; and E is the Young’s modulus. Suppose the beam thickness is 20 µm, and the width is 80 µm, the calculated silicon cantilever beam length should be 1454 µm. However, due to compromises on silicon chip size and layout constrains from the shared fabrication processes on SOI

(Silicon on Insulator) wafer, the fabricated springs are shorter than designed. The length of springs in the fabricated prototype sensor is 850 µm. For a spring with M = 20 mg, L =

850 µm, B = 80 µm, and h = 20 µm the calculated critical frequency of the system Fc is

445 Hz. The prototype sensor weighs about 25 mg and the SOI wafer used also has a 2

µm silicon dioxide on the top of the 20 µm device layer. If the difference in Young’s modulus of silicon and Silicon dioxide is neglected, the recalculated Fc is 459 Hz. Fig.8 shows the structure of the designed displacement sensor.

Figure 8 Structure of the displacement sensor. All dimensions are in micro-meter. The sensor chip is 2.2 mm by 2.2 mm square

The device is designed to be heavily damped. The damping mechanism employed in this device is squeeze thin film damping. The damping coefficient, c, can be calculated using the equation [I-10]:

LB3 1 c  h 3 (1 23/ ) 2 0 where L, B and h are the length, the width of the diaphragm and the distance of the gap between diaphragm and substrate separately. µ is the viscosity of air. ε is a constant number. The calculated damping coefficient is 1.17, and the damping ratio can be derived from the equation:

c   2 mk

The mass, m, is 30 mg and the spring constant, k, is 31.5 Nm. So the damping ratio of this displacement sensor is 0.607.

II sensor fabrication

The basic structure of the masks is shown in Fig. 9. There are three major steps to fabricate this displacement sensor, they are illustrated in Fig.10, from a) to c), in the A-A cross section view of Fig 9..The fabrication started on a four inch Silicon on Insulator

(SOI) wafer. The first step is standard photolithography process as shown in a). The second step is DRIE (Deep Reactive Ion Etching) and the etching stops when the SiO2 layer exposeed, the etching depth is 20 µm. The holes etched on the diaphragm are for the damping and releasing sacrifice layer purpose. The third step is the sacrifice layer release, which will be discussed into details later.

Figure 9 Top view of the sensor

(a)

(b)

(c)

Silicon Silicon Oxide Photo resist

Figure 10 Fabrication flow chart (A-A cross section view)

The oxide layer of the SOI wafer has a thickness of 2 µm and the device layer is 20 µm, and the size of the diaphragm is 850 µm by 850 µm. It is an issue how to release the SiO2 sacrificial layer without stiction through an efficient and relatively simple way. So we developed a HF dry etching method to release the sacrificial layer in our lab, the

37 experiment setup is illustrated in fig.8. The sensor is attached on the surface of a dummy silicon wafer with the diaphragm facing down to the 49% HF solution. A Teflon beaker is used to hold the HF solution because the HF is corrosive to the glass beaker. A tape heater is glued on the other side of the wafer with a thermal couple to monitor the temperature of the wafer. Water plays an important role in the chemical reaction between

SiO2 and HF solution; the chemical reaction procedure can be described by the following equations [I-12]

The etching of SiO2 occurs on the surface of the SiO2, so at the beginning, the gas status

HF and H2O are absorbed by the surface of SiO2.

HF (gas) ↔HF(adsorbed)

H2O(gas)↔ H2O(adsorbed)

2HF (adsorbed) + H2O (adsorbed) → HF2-(adsorbed) + H2OH+ (adsorbed)

SiO2 (solid) +2HF2-(adsorbed) + 2H2OH+ (adsorbed) →SiF4 (adsorbed) +4H2O (adsorbed)

SiF4 (adsorbed) ↔SiF4 (gas)

H2O (adsorbed) ↔ H2O (gas)

Then the absorbed HF and water mixed to start the reaction. Solid state silicon oxide reacted with acid and generated SiF4 and water which are absorbed on the surface of silicon oxide. These absorbed reaction products then turn into gas status by heating. Too much water will cause the reaction too fast and difficult to control, and the most important is the stiction because of the water vapor condensation. However, the reaction may not happen or in a very slow way if there is not enough water presents in the reaction.

The temperature is the most critical parameter to control in this experiment. Based on our experiment results, a temperature difference of 7.5 OC under our experiment condition

38 gives best etching results. The etching rate is proximally 0.36 µm / minute. In Fig 10, two pictures of the same sensor were taken under the Infrared light source. Fig. 10 a) shows the sensor before HF etching, while b) is after HF etching. The dark grey parts in the after etching picture are unreleased SiO2.

Figure 11 Dry etch setup illustration

Diaphragm Beam serves as spring

Anchor a)

Figure 12 Infrared camera photos of the device surface

a). before HF dry etching; b). after HF dry etching

Chapter 4

Sensor test setup and results The sensors were measured using the same PZT stimulation unit, and the Laser Doppler

Vibrometer (LDV) displacement measurement equipment, as described in Ping Huang’s

Master thesis [I-8]. The PZT (p-810.10 from PI Company) is used to simulate the human umbo vibration. The relation between displacement of the PZT and driving voltage under different frequency was calibrated by the LDV. The PZT could perform accurately the same way as the human umbo when excited by the incoming sound. The LDV is also used to measure the sensor and the substrate displacements, the x, and y in the theoretical model. In the following experiments, The PZT vibration amplitude is set at a constant umbo displacement level, corresponding to 97 dB SPL sound input at 1 kHz.

The measurement set-up is illustrated in Fig. 13. The sensor chip is mounted on a PCB substrate which is used to adjust the mass, M, so as to adjust the resonant frequency of the system. Due to the manipulation difficulties, instead of vertically mounted, the sensor is horizontally attached to the vibration source via a metal coupler with a thin layer of commercial hard adhesive. The whole unit is attached to a small piece of very soft foam with a very low spring constant (compared with the K of the sensor), to maintain the position of the sensor which is not shown in Fig.13. The constant gravity effect is neglected in the analysis of data. The sensor capacitance output is fed through a low noise capacitance-to-voltage converting interface circuit to convert capacitance change into voltage and the output is measured by a signal spectrum analyzer. The interface circuit is

40 connected to the sensor through 1 mil diameter flexible gold wires. The wires are long enough to minimize the strain interference to the sensor diaphragm and the spring.

Figure 13 Sketched illustration of lab test setup of the displacement sensor

Figure 14 Photos of lab test setup

Fig.14 shows the photos of the lab test setup. The optical microscope is used for the alignment purpose. In the zoom in picture, the sensor on the right side is pushed by the manipulator to the tungsten pin attached on the PZT (the pin is perpendicular to the

41 sensor mount on the soft form). The minimum screw step of the manipulator is 10 µm, but the fabricated gap between diaphragm and substrate is only 2 µm, so the manipulation process should be very carefully adjusted within a range of 0.5 µm to avoid signal distortion (If the maximum signal amplitude is “A µm” the gap should be adjusted in the range between (2-A) µm and A µm to avoid clipping of signal).

The sensitivity of the sensor module is limited by the total noise of the module. The capacitive sensor itself has very low noise and does not consume any real average power.

In order to satisfy the system sensitivity, a low-noise interface circuit is used. The circuit noise is usually specified by the input-referred noise power spectral density. For the low- noise interface circuit that is needed, the interface electronics designed for the high- performance capacitive strain sensor by this group were used [I-12]. The interface circuit has a resolution, in capacitance change, of 250 aF over a bandwidth of dc to 10 kHz, and an input-referred voltage noise power spectral density of 5 nV/(Hz)1/2. The architecture of the interface electronics is shown in Fig. 16. The displacement sensor and a fixed capacitor are used as the differential input capacitances in the “MEMS sensor” block, in

Fig.16. They are driven by a 1MHz clock signal with 3-V amplitude and are interfaced by a differential charge amplifier, which converts the sensor capacitance change to an output voltage. The high clock frequency is chosen to modulate the sensor information away from the low-frequency noise, such as the noise of the amplifier, a critical means to achieve high sensitivity. An input common-mode feedback (ICMFB) circuit and an output common-mode feedback (OCMFB) circuit are incorporated with the charge

42 amplifier to minimize its common-mode shift caused by the driving clock; hence, suppressing any offset signal due to the parasitic capacitance mismatch and drift over time. The charge amplifier output is then mixed by the same clock signal and low-pass filtered to obtain an output voltage, which represents the desired sensor information. The interface circuit is fabricated by MOSIS using 1.5 µm technology, and it consumes

1.5mA at 3V [I-14].

Figure 15 Block diagram of the low noise interface circuit

The output voltage versus input displacement characteristics of the fabricated sensor is measured using LVD in our lab, and the calibrated PZT vibration simulator [I-8]. The measured voltage output of fabricated displacement sensor, with 30 mg mass, as a function of frequency is shown in Fig.16, and the analytical response with a damping ratio of 0.6 from Chapter 2 Fig.6, is plotted in dash lines. The experiment results fit the theoretical curve within a reasonable error range in the high frequency region (before the dip on the curve). However, in the low frequency region, below the dip, the experiment results do not comply with the theoretical curve, which predicts that the output should go down with frequency without any bounce back. There is no perfect explanation for this

43 phenomenon in the low frequency region yet. Several possible reasons are proposed here for further discussion, I) The direct coupling from PZT vibration source to sensor through the air and the test setup at low frequencies. II) Due to the non-perfect setup of the test fixtures, instead of primary vibration mode, higher order mode of vibrations may dominate in the low frequency region. III) The sensor is attached to soft foam (with a very low spring constant) during the test. In the low frequency region, the diaphragm and mass are supposed to move together without relative displacement. The existence of the foam may change this fact, and cause an additional capacitance change in the low frequency range.

Figure 16 Voltage output of prototype displamcement sensor with 30 mg mass and comparison with the theoritical curve ( dash line )

From the plotted curve, the maximum signal to noise ratio (SNR) of this device reaches

74 dB at 1 kHz. As the frequency increases, the SNR decreases due to the frequency

44 bandwidth limitation of the interface circuit. This prototype sensor has a critical ffrequency, Fc, around 480 Hz, which is higher than desired value of 200 Hz for reasons, and it was explained in the previous chapters.

The displacement of thhe sensor substrate, M, of another sensor unit with same mass under same conditions was measured and is shown in Fig. 17, where the response near the critical frequency, ωc and the decrease in amplitude beyond ωc are clearly illustrated as predicted.

In order to show that Fc can be designed acccording to equation, f= (1/2π) (k/m) 1/2, the device mass of the same sensor was changed from 30 mg to 125 mg by adding a weight on the substrate. The measured results on the M=125 mg sensor showed that the cut off frequency was shifted from 480 Hz to below 200 Hz. Both of these two sensors with different masses have a near flat frequency response from 800 to 8 kHz with variation less than 6 dB.

ωc

Figuure 17 Measured displacement x of the mass M

As shown in Fig.16, for the displacement sensor with 30 mg mass, the signal to noise voltage ratio (signal analyzer bandwidth is 100 Hz) is 74 dB, at 1 kHz and 97 dB SPL input. Assume the uncertainty of the measurement is within 3 dB, and then a signal noise ratio of 71 dB can be achieved with 100 Hz signal bandwidth. Furthermore, it we assumed that a minimum S/N ratio of 6 dB is needed in hearing sound, a minimum detectable input sound level of 32 dB SPL can be achieved around 1 kHz with 100 Hz channel bandwidth with the displacement sensor.

Fig.18 compares the performances and minimum detectable Sound Pressure Level of three different approaches of acoustic /displacement sensor studies in our lab. The accelerometer [I-8] can be a very sensitive device between the frequency range from 1.5 kHz to 3.5 kHz, but in the low and high frequency band, the sensitivity drops greatly The spring coupled displacement sensor [I-9] has an almost flat minimum detectable SPL over the whole interested bandwidth. However, it suffers from a low sensitivity, which may not be able to meet the minimum sensitivity requirement as an implantable acoustic device. The direct coupled displacement sensor combines the advantages of these two approaches and can achieve a high sensitivity near 1 kHz region. In the low frequency region, the sensitivity decrease to a very low level and this weakness can be overcome by moving the critical frequency down to lower than 200 Hz.

Figure 18 Comparison of minimum detectable SPL of three

Chapter 5

Discussion and further sensor design

The feasibility of MEMS displacement sensor as an acoustic sensing device ffor totally implantable hearing aid system is confirmed in principle from the previous chapters.

However, there are problems of the prototype sensors in the sensor structure that it is not implantable. A prototype implantable displacement sensor is proposed based on our research work. This sensor is expected to overcome most of the problems in the other sensors studied previously. The cross-section view of the sensor is plotted in Fig.19 a).

The 3D cross-section view is plotted in Fig.19 b). First, the prototype sensor itself is a sealed system, which avoids the potential contaminations such as dusts and water, and provides convince for packaging also. Second, the fabricated on chip coupler greatly reduces the setup difficulties in the lab test or surgery. Third, the stopper eliminates the possible over range displacement of the diaphragm during the handling of the sensor. The set-up diagram of the sensor implanted on umbo is shown in Fig.19 c).

19-a)

19-b)

Malleus Incus (Detachable)

Stapes (Detachable) Spring made of thin diaphragm

Bio-compatible coating

Interface electronics

Umbo Ear drum Capacitive vibration (Tympanic membrane) Malleus Top plate sensor Sensor & interface circuit module Sensoor substrate (1) (Bottom electrode)

Medical adhesive gap Protective coating Middle Ear Microphone Sensor Module (Including sensor & interface circuit) (2)

19‐c)

Figure 19 a) cross-section view of cthe new design; b) 3-D view of the new design; c) implanteThde sensor setup on umbo.

The Suggested fabrication process of the protype sensor is shownin Fig.20.

Serving as diaphragm later

Silicon bottom structure

Isolation layer (SiO2)

Bottom structure substrate

(a)

Silicon top structure

Isolation layer (SiO2) + Fusion bonding Serving as spring later Silicon layer Isolation layer (SiO2)

Silicon bottom structure

(b) Displacement coupler

Silicon top structure

Isolation layer (SiO2) Silicon layer Isolation layer (SiO2) Silicon bottom structure

Protection stop (c)

Silicon top structure

Isolation layer (SiO2) Silicon layer Isolation layer (SiO2) Silicon bottom structure

Silicon bottom structure

Isolation layer (SiO2)

Bottom structure substrate

(d)

Sensor handle

Isolation layer (SiO2) Silicon layer Isolation layer (SiO2) Silicon bottom structure

Silicon bottom structure

Isolation layer (SiO2)

Bottom structure substrate

(e)

Figure 20 Suggested fabrication process of the protype sensor

The proposed prototype capacitive displacement sensor can be fabricated using micromachining fabrication process. Fig.20 illustrates the major fabrication steps. The process begins with a commercial 4-inch diameter silicon-on-insulator (SOI) wafer with a 30-m thick silicon structure, which serves as the sensor’s top plate structure later. The wafer starts with the standard photolithography process and followed by a deep reaction ion etch (DRIE) to generate the top plate structure, as shown in Fig. 20.a).

The second wafer is a special 5-layer wafer, fabricated by fusion bonding a SOI wafer to a bare silicon wafer coated with 1.5-m thermal grown silicon oxide (SiO2), as shown in

Fig. 20.b). The wafer will under two etch processes to generate the protection stop structure and displacement coupler structure, as shown in Fig.20.c).

Then the two wafers will be bonded together using fusion bonding technique. Another etching process will generate the handle structure, which will be epoxy glued to the umbo in application. Finally, the top plate structure is released by etching away the underneath sacrificial silicon oxide layer using HF vapor release process.

The micro-fabricated device is a totally silicon-sealed structure with the movable parts protected by the stoppers. Therefore the device can stand high vibration introduced by human body during sports or other events. A bio-compatible material coating can be applied to achieve the device bio-compatibility.

Conclusion

The thesis Part I verified the feasibility of the simple umbo mounted MEMS capacitive displacement sensor as implantable microphone for totally implantable hearing systems.

This displacement sensor with high pass characteristics can suppress very loud low frequency sounds, and can withstand sudden mechanical shock, such as the body movement or change of atmospheric pressure. A design of prototype sensor is also presented. The new design included considerations of a practical microphone for the totally implantable hearing systems.

Appendix I

Comparison of stable base spring mass second order system with moving base spring mass second order system discussed above

F sin ωt

x M

C k

Ground

This figure shows a simplified model of stable base second order spring mass system.

The mass M is stimulated by a force F=F sin ωt. The spring constant is k and damping factor is c. The absolute displacement of mass is x. To study the steady state response of the system, we start from writing force balance equation of mass based on Newton’s second law. The acceleration of the mass is a.

Ftotal=ma

dx d2 x Fkxc  m dt dt 2 dx2 dx mckxFsint (5) dt2 dt

λt1 λt2 This equation has two parts of solutions, transient solution xc(t)=C1e +C2e and steady state solution, xp(t)=X sin(ωt+ϕ), where C1 C2 λ1 λ2 are integration constants. We are going to calculate the steady state solution. Equation (5) has a same format with equation

(3) calculated above, which is

dx2 dx m2  c kx kA sin t  cA cos t (3) dt dt

Both the units of right side of equation (3) and (5) are newton, so they are forces. The difference is that, for equation (5), the force is independent with the system characteristic and vibration frequency. In equation (3), the total equivalent force is a function of k,c, and frequency of external excitation.

Follow the procedure of solving previous problem, the steady state response of mass M can be written as

xp (t) Xsin( t ) = Xsin sin t  X cos cos t if P=Xsin Q= Xcos then xpp (t) Psin t Qcos t,substitute x (t) into equation (5) F P(km)sin(t)2 (k m22 ) (c ) 2 cF Qsin(t)  (k m22 ) (c ) 2 F x (t)[(km)sintccost]2 p (k m22 ) (c ) 2 F x(t)p  sin(t ) (k m22 ) (c ) 2

The amplitude of this displacement response is

F X  (6) (k m22 ) (c ) 2

Compare this solution with the corresponding part of moving base system

kA 1 (2 /  )2 X'=n (k m22 ) (c ) 2

Dividing the spring constant k in equation (6) F A X k mc 22 2 (1 22 ) ( ) 2 (1 r )  (2 r) kk

ω in which r is frequency ratio , A is the displacement of mass under static load F ωn So the displacement ratio is

X1  A (1 r22 ) (2 r) 2

The figure below plotted the this displacement ratio

Bibliography [I-1] Better Hearing Institute, Prevalence of hearing loss & demographics.[Online].

[I-2] W. Ko, A. Maniglia, and R. Zhang, “A preliminary son the implantable middle ear hearing aid,” in Proc. IEEE 9th Annu. Conf. Eng. Med. Biol., 1987, p. 1890.

[I-3] A. Maniglia, W. Ko, and M. Rosenbaum, “A contactless electromagnetic implantable middle ear device of the ossicular stimulating type,” Ear Nose Throat J., vol. 73, no. 2, p. 78, 1994.

[I-4] H. Zenner,M. Maassen, R. Lehner, J. Baumann, and H. Leysieffer, “An implantable hearing aid for inner ear hearing loss: Short-term implantation of microphone and transducer,” Otolaryng. Head Neck Surgery, vol. 45, no. 10, pp. 872–880, Oct. 1997.

[I-5] A. Vujanic, R. Pavelka, N. Adamovic, C. Kment, S. Mitic, W. Brenner, and G. Popovic, “Development of a totally implantable hearing aid,” in Proc. 23rd Int. Conf. Microelectronics, Yugoslavia, May 2002, vol. 1,NIS.

[I-6] A. Maniglia, H. Abbass, T. Azar, M. Kane, P. Amantia, Garverick, W. Ko, W. Frenz, and T. Falk, “The middle ear bioelectronic microphone for a totally implantable cochlear hearing device for profound and total hearing loss,” Amer. J. Otol., vol. 20, pp. 602–611, 1999.

[I-7] D. Chen, D. Backous, M. Arriaga, R. Garvin, D. Kobylek, T. Littman, S. Walgre, and L. David, “A totally implantable middle ear device for sensorineural hearing loss,” Otolaryng. Head Neck Surgery, vol. 131, no. 6, pp. 904–916, 2004.

[I-8] P. Huang,” A Laboratory Study on a Capacitive Displacement Sensor as an Implant Microphone in Totally Cochlear Implant Hearing Aid Systems“ M. Sc, dissertation, Case Western Reserve Univ., Cleveland, OH, 2007.

[I-9] M. Zucher, “Development of a MEMS middle ear acoustic sensor for a fully implantable cochlear prosthesis,” M.Sc. dissertation, Case Western Reserve Univ., Cleveland, OH, 2006.

[I-10] Bao Min Hang, “Analysis and Design Principles of MEMS Devices” Elsevier Science; First edition. June, 2005

[I-11] W. H. Ko, R. Zhang, P. Huang, J. Guo, X. Ye, , D. J. Young, and C. A. Megerian, Studies of MEMS Acoustic Sensor as Implantable Microphone for Totally Implantable Hearing Aid Systems, IEEE Transaction on Biomedical Circuits and Systems, Vol.3, No 5, (2009), p.277-285

[I-12] Lee.Y.-I.;P.K.-H.;L.J.; L.C.-S.; H. Jounyoo; K. C.-J.;Y.Y.-S.” Dry lease for surface micromachining with HF vapor-phase etching “J.of microelectromechanical systems, 1997,vol.6,no3,pp.226-233 [I-13] M. Suster,” High-temperature MEMS wireless sensing and telemetry “Ph.D dissertation, Case Western Reserve University, Cleveland, OH, 2010

[I-14] M. Suster, N. Chaimanonart, J. Guo, W. H. Ko, and D. Young, “Remote-powered high-performance strain sensing microsystem,” in Proc. IEEE Techn. Dig., Int. Conf. MEMS, Miami, FL, 2005, pp. 255–258.

Part II.

Non-hermetic chronic micro-packaging study based on totally implantable telemetry system

Chapter 1

Introduction and Background

Introduction

Packaging is an important, nevertheless easily being ignored, and technology for modern microelectronic and micro machine systems. In most cases, the packaging determines the life time of the electronic system. This thesis, built upon many previous works in our research group [II-1,II-2,II-3,II-4,II-5,II-6], presents a novel non hermetic multilayer, multi-materials packaging technology for micro-systems, spectrally for implantable biomedical devices/systems, aimed for a life time of 0.2 to 2 years and longer. Although this study is limited to the micro-package of biomedical MEMS devices and associated micro-fabricated systems for life science research aimed to have a short to medium life time, the theory and the technology developed can be applied to industrial or electronic packages to operate under harsh environments.

Because of the small size and harsh working environment, materials can be used for implantable devices are very limited, furthermore, there is not such a single material is able to fulfill all the requirements for the packaging of implantable biomedical micro- systems. Therefore multi-materials are used to compensate each other’s imperfection as a packaging material.

The theory of multilayer coating is developed through failure analysis and literature study and testified by additional experiments. According to the theory and experimental results, multilayer coating provided a much better water resistant ability than single layer coating of the same material at same thickness. The main reason is that multilayer coating could significantly reduce probability of having through-package-defects. A micro-power telemetry unit is packaged using this method to evaluate the performance of this technology. This telemetry unit can sense pressure, humidity, and bio-potentials, for implantable biomedical micro-systems used in research, diagnosis, and therapy. The telemetry device consumes less than 1.0 μA at 3 volts, with a volume smaller than 0.15 cm3 including Lithium rechargeable battery, and has RF links to charge Li battery and to receive external on-off commands. The device output pulse width changes with a capacitive pressure sensor; and the pulse period changes with an external R as the humidity sensor or the bio-potential signals in series with R. The first model of telemetry units, packaged with silicon and epoxy only, lasted > 70 days in 40OC saline solution.

The telemetry unit packaged with 2 to 5 µm parylene C, silicon and epoxy are tested in

85 0C saline and lasted 14 days without failure, the estimated life time in 40oC saline would be longer than 896 days or two years.

Background

As the development of microelectronics and micromachining technology, the size of the traditional relatively large size medical devices are greatly shrunk down, and many of these devices can be totally implanted in human body. These implanted biomedical devices include Cardiac Pacemakers, Cardiac Resynchronization Therapy Devices,

Neurological Stimulators, Ophthalmic Implants and Urological Implants and so on.

However, the success of the microelectronics and transducer fabrication technology is necessary but not the only condition to make the implantable devices possible. Packaging of the implantable devices is a crucial technique to ensure the implanted circuits and sensors working properly, as well as to avoid the tissue reaction with the implanted systems, or to make the devices biocompatible. The packaging technology for biomedical

Microsystems has been studied in some universities and research institutes, such as Case

Western Reserve University, University of Michigan, California institute of Technology,

National Institute of Health, University of California, Berkeley and University of

California, Los Angeles. University of Michigan reported a glass package which provides a projected mean time to failure (MTTF) longer than 100 years based on high temperature accelerated tests mode. However the glass could not be the perfect material for package of biomedical Microsystems, because glass is brittle and have cracking problem as well as the problems of feed-through and difficulties to have sensors that have to communication with body fluid. Based on these reported works and previous research in our lab, there are several requirements that micro-packaging should fulfill to protect the implantable systems. They are: i) mechanically strong so that it can withstand the expected force/pressure, to isolate the active devices inside the package from undue stress and strain; ii) a good vapor barriers to prevent water and other ion vapors from permeating through to the implant device/system causing electrical leakage, corrosion and other damages; at the same time to prevent any toxic vapor from permeating out from the implant device to cause inflammation and or unacceptable pathological reactions in tissue around the implant site; iii) biologically inert and tissue compatible

61 thus do not unduly alter the function/operation of the living system: such as the outer layer material of the package should have a mechanical stiffness similar to surrounding tissue, likewise the shape and surface of package should not produce large stress, strain or high temperature hot-spots on the interfacing tissues. The specific gravity of the total package should be nearly equal to that of the surrounding tissue. The package should not incorporate sharp spikes and other structure to irritate surrounding tissue to cause a thick fibro-collagenous capsule growth around the package; and iv) sterilizable and have no biological elements (virus, proteins…) exposed or leaking out of the device. Most currently used implant packages include a hermetic seal box for vapor barrier as well as mechanical protection and a silicone like material coating for tissue compatibility and does not include the other physical and chemical considerations. But a viable, robust biocompatible package for a chronic implantable system should consider all the functions mentioned above, multiple materials would be required, especially for micro-fabricated devices to be implanted on or in an organ.

Because the high humidity environment in human body, traditional non hermetic packaging, such as dip coating of silicone, only lasts randomly from several hours to less than one week after implantation. Hermetic packaging using metal, ceramic, and glass has been the most popular method for the commercialized products, but due to the reasons mentioned previously; the cost and time needed to develop a package for implantable micro-system would be very large; and the volume of package materials would be much larger than the volume of the active components of the micro-system, these types of packaging are not suitable for most small volume prototype micro-systems tests needed in biomedical research. The new micro packaging technology developed in

62 this thesis is a step toward the solution for package of short (months) and median life time (2 years) biomedical micro-system for life science research.

Objective of part II

The objective of this thesis is to study the feasibility of the chronic non-hermetic multilayer, multi-materials coating package techniques for implantable biomedical micro- systems.

Outline of part II

Chapter 1. Introduction and Background

Chapter 2. Theory of multilayer coating

Chapter 3. Test of single material coated packaging on PCB

Chapter 4. Evaluation of the multilayer packaging based on the totally implantable telemetry unit

Chapter 5. Discussion of test results and future work

References

Chapter 2

Multilayer coating theory

This chapter summarizes the multilayer micro-package theories developed by research team members in Dr. Ko’s laboratory. They were verified by experiments from this thesis.

It was recognized early in 1970 that the packaging of implantable instruments, especially the vapor barrier layer, is one of the most significant problems that hinders the acceptance of implantable instruments in biomedical research and health care. The vapor permeability of the package is the most serious issue facing implantable electronic devices from the very beginning of implantable electronic systems for telemetry or stimulation 50 years ago.

The vapor permeability of different materials was studied for many decades. A “bird’s eye view” type summary of the permeability of various materials is shown in Fig.1-a) and

Fig.1-b). The Permeability, ρ, in g/cm-sec-torr can be defined in the equation ρ =

At(δP)/m L , where m is the amount of vapor (in gram) permeate through a film, with area A (in cm2), thickness (or permeation path length) L (in cm), in t seconds, under a pressure difference δ P (in torr ). For metal and ceramic, even at the thickness from micron meter to nanometer the permeability can be 10-12 to10-14 g/cm-s-torr.

10000 PE‐LD EVA 1000 PS /day)

2 PE‐HD PP PC BOPP ETFE /m

3 COC 100 ECTFE

(cm PCTFE PVC PVDF 10 PET Rate PA 6 PAN PEN ORMOCER I 1 PVDC EVOH ORMOCER II

Transmission 0.1

2 LCP O

0.01 0.010.1 1 10 100 Water Vapor Transmission Rate (g/m2/day) b) Figure.1. (a) Permeability of polymeric materials [II-6]. (b) Effectiveness of the

packaging materials [II-7-II-10].

Up to the present, hermetic packages in bulk metal, ceramic, or glass boxes are used for chronic implants. The thickness of the package is mostly determined by the mechanical requirement rather than the vapor permeabilitty. The micro-package research is aimed to develop packaging techniques to reduce the volume, size, and weight of the package to be

65 comparable or smaller than that of the active part of the micro-system; and to fabricate the packages with micro or nano-fabrication technology at reduced time and cost for biomedical implantable micro-systems used in research and health care.

Theory I. The Non-hermetic Package

The hermeticity of package requirement is based on when there is an air or vacuum space inside the package box. However, when there is no box, no void or air space in package, the concern on protecting the device from water or ion vapor would be to have a survival time longer than the desired device operation life time. And the survival time is the time for the water or other vapor permeates through the package layers and causing: reduction of the package resistivity, corrosion, and other harmful effects of the packaged device.

Then it is not necessary to have “Hermetic” package and it is possible to have non- hermetic package when: 1). There is good bonding and no void or air space between the package layer and the active system; 2). The base board of the active system and the active components are clean without ionic contamination and residues; and 3). the package layers maintain high resistivity when saturated with water vapor. Under these conditions the water vapor may saturate the package layers but will not reduce the resistivity of the package and will not cause low surface resistivity that cause electronic device failure.

Many materials, such as silicone, the vapor can permeate through mm thick layer in hours and days but the resistivity of the water vapor saturated silicone film would still be very high. If the bond between the active device surface and the package layer is good, there is no void for vapor to condense into water pocket when temperature is cycled and the

66 active device and base board surface has no ionic contamination to react with the water in the package material, then the device would see a high resistance sheet of packaged material and can continue to function properly. As long as the package layer presents an environment with higher resistivity than required for device to function properly, even the package is not “hermetic” the packaged implantable device can function normally.

The minute ionic contaminate on circuit board is very difficult to eliminate. The results from literature and our early works indicated that the silicone dip coated package devices usually failed randomly in hours and days. Water vapor can penetrate the packaging through two major routes, one is through the coating layers, and the other is through the defects. The penetration through defects is obviously much faster. The defects include air bubbles, ions, dust particles and foreign material fractions. It is observed in our previous works that the life time of the package is limited by the defects and uneven coating.

When careful coating technique was developed to coat silicone in three or four thin layers, each with 50 to 100µm thick, the package life time with silicone alone extended beyond

40 days [II-11].

Theory II: The multiple layers coating reduces defects caused failure

There are two parts in theory II, they are: A). Multiple coating of the same material in thin layers reduce the failure caused by the defects; and B). Multilayer of different materials reduces the failure by Defect-decoupling effect

A. Multiple coating of the same material in thin layers, with the same total thickness of the material will reduce the failure caused by the defects.

In an amorphous polymer body, the “universal” free volume fraction value iss 2.5% [II-

12]. So the total ratio of defects to body volume in a coating layer is at least 2.5%. It is believed that multilayer coating has the advantage of reducing defects overlapping probability. A theoretical reasoning is developed to support this expectation as follows:

Third layer

Second layer Defects First layer

Substrate

Figure 2. Schematic of defects in a multilayer coating

Fig. 2 shows the schematic of defects in a multilayer coating, which has three layers on top of the device. The defects are randomly distributed on each layer. When the defects on different layers overlap to the one on other layers, they formed “through defects”, which are the easiest way for water vapor to penetrate and cause early failures. The life time improvement of a multilayer packaging can be estimated by calculating the reduced probability of through defects. The density of through defects is a function of defects concentration of each layer, layer area and number of layyers. If the probability to have a defect, such as an air bubble, on a coated film with area A, thickness d, is γ , the size of the bubble would be smaller than or equal to d. The probability to have another bubble

2 2 near the same location within an area (nd) , on the second layer of film would be γ2 =γ

{(nd)2 /A}. For 3 layers film, the probability to have a bubble at the same location within

2 3 2 2 2 area (nd) would be γ3 =γ {(n/d) /A} . If A = 50 mm , γ = 0.01, d = 0.05 mm, n =2,

2 -2 2 -11 (nd) = 10 mm , then γ3 =4x10 which is very small. Therefore the chance to have

68 bubble like through-defect failures would be greatly reduced for device coated with several thin layers of the same material than one coat with same total thickness.

Recent publications from Russia[II-13,II-14], gave more rigorous analysis on probability of defects when spray coating Teflon films on large area. They assumed: (1) dimensions of defects are much smaller than device dimensions; (2) defects sides are orthogonal to the coating surface and (3) All defects are circular shaped with a radius R; the probability of having a defect in single coating over the area is n, then, Pm, the probability of the formation of a through-defect of coatings is:

1 Pn(4R) m2m1  (1) m Q m1 where Q is the ratio of defect area to the coating area, n is the number of defects on a single layer, m is the coating layers, and R is the diameter of the circular defect. For 6- layer spray coated material, the probability of having a through-defect is 3.18x10-11. This confirmed the hypotheses proposed for this project.

B. Multilayer of different materials reduces the failure by the Defect-decoupling

effect.

The barrier performance by single layer is limited by defects on the single film; however, with combining dissimilar layers of materials, both simulation and experimental data showed the improvement of water vapor transmission rate of combined film is much lower than the addition of two single films acts alone [II-15]. The results are shown in

Figure 3.

1000

100

10 PC/par/ PC/par/ PC/par/ SiNx/par/ SiNx SiNx/par /day) SiNx 2 1 PC PC/ parylene (cc/m 0.1

Oxygen Tramission Rate 0.01

0.001

When using ceramic or any brittle material film as vapor barrier, the development of high stress or cracks under mechanical loading and mismatches in temperature expansion coefficients are major concerns. Cracks occur when the maximum tensile stress exceeds the stress limit of the material. To estimate the stress of a film on a substrate, the Stoney formula is usually used [II-16,II-17]. When a film is deposited on a substrate as shown in

Fig .4, the film stress, f can be expressed as equation (2).

Figure 4. The structure of a film on a substrate

3 Eds  f  (2) 2 d s 6(1 ) R df (1  ) d f Where R is the radius of curvature when the structure is bended, E and  are Young modulus and Poisson‘s ratio of the substrate, and ds and df are the thickness of the substrate and the film respectively. For a given R, the stress,f, is proportional to E and

3 2 ds of the substrate and inversely proportional to (df ) of the film. In order to reduce the stress, one should reduce film thickness, substrate thickness and substrate Young’s modulus (Stiffness). When df is reduced from 1 mm to 1 nm the stress will be reduced by a factor of 106. Crack can be avoided most of the times.

Furthermore, if the ceramic thin film is deposited on a flexible thin substrate, or sub-layer, then df and ds are small, and E is also small. Then the film stress would be greatly reduced, and no cracking would occur in the ceramic vapor barrier films under bending.

For a multiple layer packages the ceramic films are interlaced with thin films of flexible polymeric material, therefore the stress of the ceramic films would be minimized to prevent cracks from developing. If layers of very thin ceramic films is interlaced with thin flexible materials to form a multi-layers package the stress in each layer of the overall structure is greatly reduced. In this way, crack can be prevented even under repeated bending of the total structure, and the mechanical flexibility of the total package can be achieved.

A flexible material with a small Young’s modulus could also adsorb the inter layer stress and does not transfer stress from the bottom substrate surface to the front surface across

71 the film. Therefore the temperature coefficient difference between the ceramic films and substrate, or a lower coating that houses the active device, would not cause a stress build up in the multi-material package layers when the package is temperature cycled.

Temperature compensation also can be accomplished by selecting interlace materials.

From these aforementioned reasoning and analyses, the proposed approach to interlace the ceramic vapor barrier thin films with thin films of flexible and compliable materials such as parylene, or some porous oxides, to absorb interface strain is expected to yield long life time, thin vapor barrier micropackage of implantable devices.

Based on these reasoning, the proposed micropackage is to use (a) a tissue compatible material such as silicone as the outer layer with the proper shape and surface structure to match the softness of the tissue around the implant device; (b) parylene C as a vapor barrier, interlaced with ceramic thin films if possible; and (c) an epoxy filled with ceramic grids as the inner mechanical protection layer; to achieve long term (months to years) micropackage of implantable devices/systems.

For the telemetry unit we used to evaluate the multi-layer micropackage techniques of this research, three layers were coated on the device. The first layer is FP 4450, which supplies mechanical support on the device, and then one or several conformal layers of

Parylene-C is deposited to serve as a water vapor barrier, the outside layer is MDX 4210 silicone, which makes the packaged micro-system biocompatible with surrounding tissues.

Chapter 3

Single packaging material coating experiments

Silicone MDX 4210, FP 4450 and Parylene C were studied as first set of candidate materials. In order to evaluate the materials and coating techniques, these three materials were coated on the printed circuit test boards with comb shaped electrodes with different spacing between two sets of electrodes. Then the coated test-boards are exposed to the working environment or accelerated test conditions. The surface resistance between the set of comb electrodes after coating and exposed to the harsh environment, such as saline solution at set temperature for biomedical systems, is monitored, when the resistance drops from 2 G-Ohms to below 120 M Ohm, the coating is considered failed, and the test-life-time is recorded. The test-life-time reflects the materials’ ability to withstand the test environment. For biomedical systems, the H2O and ions vapor penetration permeability is the most important factor that determine the life time of the implantable system. Besides the inter-electrode resistance, the body resistances of these package materials before and after saline bath tests are measured and recorded

3.1 Test set-ups for life time evaluation of packaging material in 40 and 85 oC saline solutions.

The single packaging material characteristics were studied under two experimental conditions-- body temperature mode and high temperature accelerated mode. The body temperature mode refers to the experiments under one atmosphere pressure, immerged in

40OC saline solution and 100% relative humidity (RH). The accelerated mode keeps the

73 pressure and humidity same as the normal condition, but boosts the saline solution temperature to 85OC. The accelerated test would significantly reduce the material mean time to failure (MTTF). The accelerated ratio of time-to-failure between the two test modes can be calculated by the following formula (3)

n T RHE11   a n exp   (3) TRnaHkTT  an 

Where, RH is Relative Humidity, ∆E, the reaction energy, is assumed to be the 0.9 eV; K,

-5 the Boltzman Constant, is 8.62×10 eV; n, an experimental number, is 3.0. Ta and Tn are temperatures in Kelvin. According to the calculation result, the failure of the device under accelerated mode test is supposed to be 64 times faster than the normal test, or the mean time to failure (MTTF) of accelerated test is one sixty fourth of the normal test.

This is a theoretical expectation based on statistic data. However, because of the different mechanism of failure reason and different chemical-electrical reaction process due to the various materials, this number may not be accurate enough for every cases and every materials. The high temperature may also cause some additional failure modes which will not happen under low temperatures. A calibration experiment of the acceleration ratio between normal test and accelerated test is conducted and will be discussed later in this chapter.

The normal mode test is set up in two beakers as shown in Fig. 5 The small beaker is filled with saline solution and sits in the larger beaker which contains some de-ionized water. The purpose of this set up is to reduce the vaporization of the saline solution from the inner beaker. The hot plate (model) has a close loop temperature control function, the

74 feedback temperature signal is supplied a thermal couple which will make sure the temperature is kept exactly at 40 OC. The Teflon tape between the covering cap and the beakers improve the sealing of the test system. Fig. 6 shows the accelerated test setup. All the devices were tested in the constant flow box (model). There is no feedback control function on this constant flow box, so an external thermal meter is used to monitor the saline temperature inside the stainless steel box.

Figure 5 Normal-Mode test setup

Figure 6. Accelerated-Mode test setup

3.2 Experimental procedure and results.

The first step in evaluation experiments is to prepare the Printed circuit board (PCB) for test. The PCBs are all designed and fabricated in our research lab. The size and feature of types of a1 and a2 PCBs with comb finger electrodes are as illustrated in Fig.7

1mm 4 mm Connecting to ohm‐meter

(a) Type a1 PCB

2 mm 6 mm Connecting to ohm‐meter

(b) Type a2 PCB

Figure.7 Layout of the PCB for material test experiment. a) narrow PCB-a1; b) wide PCB-a2

There are two types of PCB designed, and the difference between these two designs is the board width and separation distance of the electrodes to the edges. The type-a1 board has a margin of 1 mm and the type-a2 board has an edge margin of 2 mm. For both type a1and a2 boards, the comb fingers have 1 mm separation from each other. The different margin on the testing devices is aimed to evaluate the effect of water vapor invasion from the

PCB edge due to poor bonding of the coating layer to the PCB.

The preparing of coating materials is an important step which has to be handled carefully to avoid invasion of the dust, ion and other contaminants. The Dow Corning medical grade silicone MDX 4210 includes two parts, MDX-4 4210 and cure agent. The HYSOL black epoxy FP 4450 and dam materials FP 4451 are preserved in the refrigerator with

76 the storage temperature set to -40OC. The eppoxy is taken out from the refrigerator upon the immediaate usage. The surface cleaning of the PCB is the most important factor that determines the life time of the testing. Because based on the non-hermetic multilayer packaging theory, and our test results of MDX 4210 silicone, the surface resistance of the water vapor saturated silicone packaging layer, is still very high ( > 108 Ohm).

The type a1 and type a2 PCBs are used for the evaluation of MDX and ENCAP materials.

The cleaned PCB, with resistance between electrodes greater than 1000 M-ohms, is carefully coated over with multilayers of the material following the developed processes and then is cured. The finished PCB, then, is immerged in 40oC saline solution until the inter electrode resistannce fall below 100 M-ohms. The time in the saline is the expected life time of the PCB coated with the test material.

3.2.1 Failure mode observation

The locations of these failed points are very small and hard to find even under microscope. An electrolysis meethod was developed. A DC voltage with a series resistance is applied to the two leads (a, b or a, saline or b and saline) that developed low resistance as shown in Fig.8.

Figure 8. The principle of electrolysis method used to locate the leak point on the test board.

A low current (several µA) will flow through between the electrodes and saline.

Positively charged ions Na+ move towards the cathode, while negatively charged ions Cl- move toward the anode. The general reactive equation is

2NaCl+2H2O==electrolysis==2NaOH+H2↑+Cl2↑

The hydrogen bubbles can be observed at cathode leaking point. Chlorine is absorbed by sodium hydroxide. NaOH is strong electrolyte，it became Na+ and (OH) － in water:

NaOH ＝Na+ ＋ (OH) －

+ －－ + While at the anode leaking point, copper is oxidized into Cu2 , [Cu 2e = Cu2 ], resulted in Cu (OH)2, Cu(OH)2 is a blue flocculent precipitate. So we can observe some blue substance at the anode leaking points. With this method, small low resistance failure locations will show up as discoloring spots on the electrode. However, the coating faults are very small the leakage resistances are of the order of 10 to 100 Meg Ohms. With DC

30 voltages applied, it would take several hours to make the leaking points observable.

3.2.2 The evaluation of MDX4210 silicone

SILASTIC MDX4-4210 BioMedical Grade Elastomer is designed for medical device encapsulating and mold-making application where cure is at room temperature or slightly elevated temperatures. It has, for example, been used as a flexible mold to facilitate the encapsulation of electronic components of biomedical devices. The MDX 4210 can be used for general prototyping, molding and fabrication of medical device components, and also can also be used as a drug matrix for controlled release drug delivery systems.

The MDX 4210 is coated on both a1 and a2 types of PCBs. For the a1 type PCB, there were four devices (packaged in Sept. 2008) tested, and the test results are summarized in

Sample Number Coated Material Coated Thickness(µm) TBF*(hour/day)

#6 MDX 4-4210 91.5 482/20

#9 MDX 4-4210 150.5 456/19

#10 MDX 4-4210 277.5 146/6

#7 MDX 4-4210 381.0 914/38

Table 1. Test results of a1 PCB coated with multi-layer of Silicone in 9-2008

The a2 type PCBs (packaged in May 2009) test results were shown in Table 2. The a2 type has an observable improvement on the life time compared with the a1 type PCBs.

The longest one lasts 53 days.

Sample----Mab Survived days Description (ab is thickness)

M63 (2 37 Soldering area protected with ENCAP layers)

M68 (2 2 Soldering area protected with ENCAP---Soldering layers) area/wire fail

M76 (2 39 Soldering area protected with ENCAP layers)

M104 (2 7 Soldering area protected with ENCAP---Soldering layers) area/wire fail

M114 (2 43 Soldering area protected with ENCAP layers)

M118 (2 0 Soldering area protected with ENCAP---Soldering layers) area/wire fail

M127 (2 22 Soldering area protected with ENCAP---Soldering layers) area/wire fail

M143 (2 4 Soldering area protected with ENCAP---Soldering layers) area/wire fail

M154 (2 layers) 53 Soldering area protected with ENCAP

M204 (3 layers) 35 Soldering area protected with ENCAP

M219 (3 layers) 48 Soldering area protected with ENCAP

M235 (3 layers) 42 Soldering area protected with ENCAP

M256 (3 layers) 23 Soldering area protected with ENCAP---Soldering

area/wire fail

M270 (3 layers) 42 Soldering area protected with ENCAP

M345 (3 layers) 42 Soldering area protected with ENCAP

Table 2 Test results of a2 PCB coated with multi-layer of Silicone coated in May 2009

There are 15 a2 type devices were put to test, and 12 devices of them had been observed failure points in failure mode test. In the 12 devices, three devices have failure points

80 occurred on the PCB surface, they are M235, M256 and M270, the average thickness is

254 µm and the average life period is 35 days. 8 devices have failure points observed on the wires, the average thickness is 175 µm, and the average life period is 32 days. The

M118 failed immediately. M68 lasted two days, M143 lasted for four days and M104 lasted for seven days. These four devices are not included for average life period calculation. Three of them did not show any obvious failure points after 4 days test (test current is ~10 µA). From the test results, if, the average life time of silicon coated devices is longer than 35 days. When the coating thickness is over 100 µm, there is no obvious life time improvement with increased thickness of silicone. The impression from these data is that beyond 100 µm thickness multilayer coated (2 or more layers) of silicone would be able to protect the device for 30 days or longer.

Figure 9. Lifetime test results of silicone packaged PCB; dashed line is the

linear curve fitting of good device data, Red line circled data are devices

failed prematurely with identified failure mode.

3.2.3. FP 4450 (ENCAP)

Hysol® FP4450 is high purity, low stress liquid encapsulated with good moisture resistance and an extended working life. It is designed for protection of bare semiconductor devices. Pressure pot performance on live devices is up to 500 hours with no failures, depending upon device and package type. This material is designed for temperature cycling ranges up to -65°C to 150°C. Pot life or working life has been extended to approximately 3 days. This liquid epoxy exhibits relatively high flow. A cavity or a potting dam is required for flow control. Hysol® FP4450 may be suitable for bare chip protection, such as IC memory cards, chip carriers, hybrid circuits, chip-on- board, multi-chip modules, ball grid arrays and pin grid arrays. The high temperature performance and excellent resistance to chemicals, moisture and handling damage, are also its advantages.

The FP 4450 is coated on a1 and a2 types of PCBs. There are two a1 type, (packaged in

September 2008), tested, 50 µm and 330 µm. The life times in 40 OC saline solutions are

13 days and 39 days separately. The results of a1 type devices are summarized in Table 3.

Sample Number Coated Material Coated Thickness(µm) TBF*(hour/day)

#1 ENCAP 50.0 320/13

#2 ENCAP 330.0 941/39

Table 3 Test results of a1 PCB coated with Multilayer ENCAP Epoxy in 9 2008

The a2 type devices listed in Table 4 , packaged on July 8 2009, shows an improved life time compared with the a1 type devices. This phenomenon is the same as what happened

82 in the MDX 4210 experiment. The wide edge margin does have a positive effect on the device life time.

Thickness Survived days Description

E153 43 Soldering area protected with ENCAP

E203 48 Soldering area protected with ENCAP

E230 48 Soldering area protected with ENCAP

E246 93 Soldering area protected with ENCAP

Soldering area protected with ENCAP---Soldering E260 2 area/wire fail

Soldering area protected with ENCAP---Soldering E109 1 area/wire fail

Soldering area protected with ENCAP---Soldering E124 4 area/wire fail

Soldering area protected with ENCAP---Soldering E95 1 area/wire fail

Soldering area protected with ENCAP---Soldering E170 4 area/wire fail

Soldering area protected with ENCAP---Soldering E161 4 area/wire fail

Soldering area protected with ENCAP---Soldering E213 1 area/wire fail

E195 11 Soldering area protected with ENCAP---Soldering

area/wire fail

E233 32 Soldering area protected with ENCAP

o Table 4 Test Results of a2 PCB coated, with Multilayer ENCAP on 8/7/2009, in 40 C

Saline.

There are total thirteen devices tested, eight of them failed because of soldering area/wire protection problems (~71%). The other 5 devices have an average thickness of 213 µm, and average surviving time is approximately 53 days.

Table 5 listed the experiment results of the a2 type devices, packaged on 10-15- 2009, tested in 85 OC accelerated mode. Average equivalent life time converted to 40 OC is 173 days.

Thickness Survived days Description

196.5 148 Soldering area protected with ENCAP

208 244 Soldering area protected with ENCAP

291 147 Soldering area protected with ENCAP

176 16 Soldering area protected with ENCAP---Soldering

area/wire fail

303.5 10 Soldering area protected with ENCAP---Soldering

area/wire fail

281.5 10 Soldering area protected with ENCAP---Soldering

area/wire fail

185 135 Soldering area protected with ENCAP

263.5 267 Soldering area protected with ENCAP

282 140 Soldering area protected with ENCAP

323.5 132 Soldering area protected with ENCAP

Table 5 the test results of a2 PCB coated with multilayer ENCAP on10/15/2009, tested in 85oC

The test results of devices listed in Table 4 and 5 are ploted in Fig..10 a) and 10 b) with different acceleration ratio. Fig.10 a) used the acceleration ratio of 64 as calculated from the equation 1. The life times of good ENCAP coated devices tested in 40oC and 85oC separated widely. Fig.10 b) used the acceleration ratio of 20. The life time of 40oC and

85oC are closely matched. This suggested that the correct acceleration ratio for ENCAP would be around 20.

a) b)

Figure 10 Test results of ENCAP coated PCBs listed in Table 4 and 5, epresents

10-a). The life time tested in 85 oC is converted to equivalent life time at 40OC test using the accelerated ratio of 64 ; 10-b). The life time of devices tested in 85 oC is converted to equivalent life time at 40OC test using the accelerated ratio of 20

3.2.4. Parylene

Parylene is a general name for a variety of chemical vapor deposited polymers which are used as moisture barrier and high resistance electrical insulation materials. Parylene C is one of these polymers, and it has been the most popular candidate of this type of material over last 40 years because of its low cost, high performance and other advantages. It also has been widely studied by lots of research institutes and universities. In our experiment,

Kisco diX-C parylene is used. The deposition process is done by specialty coating systems PDS 2010, which is shown in Figure 11.

Figure 11 Specialty Coating Systems (SCS) PDS 2010 Deposition System

The deposition of the parylene C is completed in the room temperature, vacuum chamber.

The deposited thickness is linearly proportional to the loaded parylene C dimer weight.

This relation for the specific machine we used in Case Western Reserve University is shown in Figure 12.

Parylene C Deposition 30 y = 0.6217x ‐ 0.0896 25 R² = 0.9992 m)

 20 (

Thickness 5

0 0 1020304050 Dimer Weight (g)

Figure 12. Parylene C film thickness is linearly proportional with the parylene dimer weight

O In the first run experiment, the Parylene is coated on a2 type PCB, and tested in 40 C saline solution. The devices coated with the thickness of 25 µm are tested in 85OC saline accelerated mode. The test results are listed in Table. The life period scatters for the parylene thickness under 25 um. The failure mode analysis after the test shows that the surface contamination is the primary reason for the most of the device failure. The results are listed in Table 6.

Thickness (µm) Survived days

2 Failed immediately Soldering area protected with ENCAP

5 Failed immediately Soldering area protected with ENCAP

7 Failed immediately Soldering area protected with ENCAP

7 3 days Soldering area protected with ENCAP

9 7 days Soldering area protected with ENCAP

9 Failed immediately Soldering area protected with ENCAP

25 900 days (still testing) # Soldering area protected with ENCAP

25 900 days (still testing) # Without ENCAP protection on the

soldering area

Table 6 Life time test results of firrst run experiment with Parylene coated a2 PCBs tested in 40oC saline (#: Tested under accelerated mode, converted to normal mode life time with conversion ratio 64)

Figure 13. Failure mode observations of 2 µm parylene coated devices

The second run experiment of parylene is carried on a3 type of PCB, as shown in Figure

14. The a3 type of PCB has ten devices on one PCB, in which there are five wide figure gaps (1.25 mm) and five narrow figure gaps (0.5 mm). All these ten devices share a common ground. The resistances are measured between the ground wire, a, and the other

88 ten wires, b1 to b10. When the resistance measured is lower than 100 M Ohm, the device is considered as a failure. The experiment setup is shown in Figure 15.

b1 b2 b3 b4 b5 a b6 b7 b8 b9 b10

Figure 14. a3 type PCB test board

Figure 15. Accelerated test set-up for a3 type PCB coated with parylene C

In the second run test of parylene, all tests are in the accelerated mode, the survived days are equivalent life time converted from days in accelerated mode test (85 OC saline solutions) to days in normal mode test (40 OC saline solutions), and the conversion ration used is 64. The test duration was set to 256 equivalent days in normal mode, or 4 days in accelerated mode. There are two groups of devices tested, the thickness of parylene coating distributed from 5 µm to 25 µm. In the first group, three PCBs are coated with single layer of parylene. The thicknesses are 5 µm, 7 µm and 11 µm. Ten 5 µm single layer devices all survived after 256 days equivalent test. For the 7 um single layer devices, 80% survived for both wide and narrow finger gap devices. However, for the 11 um devices, 4 narrow gap devices and 2 wide gaps devices failed, the surviving rates are

20% and 60% separately. The test results are plotted in Figure 14-a. In the second group, there are five PCBs are coated with multilayer parylene C. Multiple depositions were conducted, and each deposition process adds 5 um parylene on top of the previous layer.

The thicknesses are 5 µm, 10 µm, 15 µm, 20 µm and 25 µm separately. The multilayer parylene coated devices shows a longer life time compared with corresponding same thickness single layer devices. The 5 µm devices are all survived. Four narrow gaps 10

µm devices and one of the wide gaps 10 µm device survived 230 days. 3 wide gaps 15

µm devices survived after 256 days (60%). 5 of the narrow gaps devices with 15 µm multilayer parylene failed immediately after immersing into saline solutions. All 15 µm and 20 µm multilayer devices survived 256 equivalent days accelerated test. The results are plotted in Figure 14-b

Comparsion of Survival Days for Single Layer Parylene

300 5 wide & 250 5 narrow

Days 4 wide & 3 wide &

200 4 narrow 1 narrow

150 Survived

100 Wider Gaps 50 Equivlaent Narrower Gaps 1 wide 2 wide 0 1 narrow 4 narrow 135791113

Parylene Thickness (micron) . . Figure 16-a. Single layer Parylene test results (with 5, 7, 11 m thickness) Survival percentage: wider gap-80%; narrower gap-67% Comparsion of Survival Days for Multiple Layer Parylene 300

250 5 wide & 4 wide 3 wide 5 wide & 5 wide & Days 5 narrow 5 narrow 5 narrow 200 1 wide & 5 narrow 150 Survived

100 Wider Gaps 2 wide &

Equivlaent 50 Narrower Gaps 5 narrow 0 0 5 10 15 20 25 30 Parylene Thickness (micron)

Figure 16-b. Multilayer Parylene test result (with 5, 5x2, 5x3, 5x4, 5x5 m thickness) Survival percentage: wider gap-92.5%; narrowergap-80%

The test results of parylene coating showed that the multiple layer coated PCB devises has 92.5% and 80% survival rate for wide gap and narrow gap PCBs; while the single layer coated PCBs survival rate are 80% and 67% for wide gap and narrow gap PCBs.

These results are consistent with test results of Silicone and ENCAP coated PCBs. They

91 clearly show that the multilayer coating yields much better results than the single layer coating.

In the third run experiment of parylene coating, all the devices are coated with single layer and multilayer 2 µm parylene. The test results are to be provided.

3.3 Calibration of life time ratio between 40 oC and 85 oC accelerated tests

The multilayer silicon coating is chosen to calibrate the accelerated test. Four devices are made in a same batch with the thickness are all in the range of 150 µm ± 20 µm. Two of them are put in 40OC saline and the other two are put in 85OC saline. All the soldering areas of the four devices are protected by ENCAP. The #1 and #2 devices with the thickness of 114 µm and 103 µm were tested in 85OC saline. #1 device lasted for 13 hours and #2 device failed immediately. The #3 and #4 devices with the thickness of 110

µm and 98 µm were tested in 40OC saline, #3 device survived for 21 days and #4 device survived for 16 days. The calibration of accelerated test ratio: 21 days * 24 hours/ 13 hours ≈ 38.8

The first calibration try of accelerated test ratio is 38.8, for silicone coating, compared with the book value of 64. Additional careful calibration with large samples should be made to find more accurate ratio value in the future.

3.4 Summary of the results and discussion.

Three materials, FP 4450, MDX 4210 and parylene C, were used as packaging materials for three types of PCBs with inter-digitated fingers in a series of Single material coating experiments to evaluate the theories proposed in Chapter 2. Even with small number of

PCBs tested on each material and each coating methods. The experimental results clearly support the proposed theories. A) All three materials are not hermetic package materials.

Using conventional coating method to coated PCBs with electrodes and circuits with silicone or epoxy the device would fail, in 40oC saline bath, from several hours to a few days. With the non-hermetic package technique, thin silicone or epoxy films coated (100 to 200 µm thick) PCBs survived more than 60 days in 40oC saline bath. The survival time is more than 10 times longer. The thin parylene (5- 25 µm thick) coated PCBs may have survival time greater than 5 months to 2 years. Therefore, the non-hermetic package is shown to be feasible B). For all experiments using all three non-hermetic materials, multi-layer coating technique is proven to be significantly better than single layer coating even all other coating techniques are the same. The early experiments on silicone and

ENCAD epoxy also showed that PCBs coated with both materials has longer survival time. These experimental results support the theory II on multi-layer coating approach.

The limited experiments in this chapter proved that the theory I--Non-hermetic micropackage is feasible and the Theory II-- Multi-layer Thin-film packaging is the correct approach for micropackage. Both theories deserve more careful study to understand the underline physics and to develop improved technology for low cost, fast turn-round packaging for implantable biomedical micro devices and systems.

In the next chapter, an implantable telemetry unit is designed and packaged using the non-hermetic packaging technique developed above to demonstrate the true correlation between the results from laboratory experiments and the functional active telemetry device in simulated body environment.

Chapter 4

Implantable telemetry unit packaged using multilayer coating techniques

Based on previous research works in our lab and other institute [II-1-II-3], we concluded that the implantable micro-package needs to :i) be biocompatible with interface tissue; ii) block vapor permeation both way; and iii) provide mechanical protection of the implantable devices; most importantly, iv) have minimal volume and weight, and is desirable to be less than the unpackaged unit. The traditional metal or ceramic packages cannot satisfy all these requirements and require large cost and time delay.

An ultralow power RF telemetry integrated circuit is designed in our lab and fabricated by MOSIS used to evaluate and demonstrate the new non-hermetic approach in micropackage. [II-11]The micro-power telemetry unit can sense pressure, humidity, and bio- potentials, for the study of micro-package techniques for implantable biomedical micro- systems used in research, diagnosis and therapy. The telemetry circuit is designed based on the 0.35 µm MOSIS technology. The block diagram working principle of micro- power telemetry device is shown in Fig. 15. The output signal of this demonstration device is a low duty cycle modulated RF pulses. The pulse period, t1, and pulse width, t2, as sketched in Fig.15, is determined by the value of off-chip resistance, R, and off-chip capacitance, C, separately. When the R and C sensors change with measured parameters

(pressure, potential or stress), the pulse period and pulse width will change. When water vapor penetrates through the package, the value of R, C will also change with humidity in a slow mono direction way. The telemetry device consumes less than 1.0 µA at 3 volts.

The lithium battery used for powering the whole system is a commercial product,

Panasonic ML-414RM. It has a 1.0 mAh capacity and theoretically can last about 1,000 hours on this telemetry circuit. The assembled telemetry unit with a volume smaller than

0.2 cm3 including lithium rechargeable battery, and will have RF links to charge Li battery and to receive external on-off commands. The first model of telemetry units, packaged with silicon and epoxy only, lasted more than 70 days in 40oC saline solution.

The RF recharging and on-off control circuits was tested, and the lifetime of second generation telemetry unit are being evaluated.

Figure 17. Sketch of output signal and circuit diagram

The telemetry IC chip is measured 0.77 mm long and 0.57 mm wide. A photograph of the fabricated chip is shown in Fig.16. This IC chip together with external components, including R, C, RF coil and the lithium battery are asseembled on a thin printed circuit board. The whole unit after assembling measures 5mm wide, 20 mm long and 2mmm thick.

The assembled telemetry unit is sketched in Fig.17.

Figure 18. Telemetry IC circuit layout

Figure 19. Sketch of assembled telemetry unit

The telemetry unit is characterized in order to evaluate and demonstrate the sensitivity of humidity and R, C, sensors of the packaged unit. The telemetry unit is able to sense humidity change, bio-potential change as well as the pressure change with a MEMS capacitive sensor.

Figure 20. Capacitance and pulse width change

The MEMS capacitive pressure sensor is connected to the telemetry circuit as the external capacitor C through wire bonding as shown in Fig.19. The capacitive sensor has a nominal capacitance of 10 pF. The sensor annd the telemetry circuit are put in a variable pressure chamber. Fig.20 plots the capacitance change versus gauge pressure and output signal pulse width versus gauge pressure. The nonlinearities of both these curves are less than 1% in the pressure range of 0 to 25 cm water.

The external components R on circuit board is working as a humidity sensor which senses the slow drift of humidity in the packaging. The pulse period t1 will decrease as the resistor R decreases, which reflects a humidity increase in the packaging. The pulse period versus resistance is plotted in Fig.21. In the range of 1.5 Mega ohms to 10 Mega ohms, the nonlinearity of the plotted curve is less than 1%.

Figure 21. Relation between pulse period and bias resistance

The circuit can also be used as a bio-potential sensor to sense small transient voltage signal, when the circuit is setup as shown in Fig.22. The pulse period changes from 50 µ seconds to 3300 micro-seconds as the input electrical voltage varies from -110 millivolts to 80 millivolts.

Figure 22. Relation between pulse period and electrical potential

The first step is to clean the telemetry unit. The dust and ion contaminations are the major issue for the surface cleaning. The ion contaminates the circuit board surface may come from a couple of different sources, such as soldering flux, dust in air and material impurities. The non-hermetic micro-packaging requires a careful cleaning process on the device surface. For the ideal condition, if there is no cavity in the package, no ion or other contaminations on the surface. The water vapor saturated packaging material keeps a high volume resistance (> 100 Mega ohms), the telemetry unit will not develop a large leakage current to cause working problems.

The assembled telemetry unit before packing is cleaned according to the following procedure. First, the PCB is cleaned in electronic grade alcohol flow; second step, rinse the PCB under DI water; thirdly, flush the PCB with electronic grade Acetone then rinse the PCB with DI water again. After the cleaning process, the telemetry unit is baked in oven for 1 hour under 80oC to drive moisture out of the circuit. The resistor, R, on telemetry circuit which senses the surface resistance change of the packaging is 20 Mega ohms. That means the surface insulation resistance should be much bigger than 20 Mega ohms to keep the circuit working properly. A 250 µm thick MDX 4210 silicone sheet was immerged in 40oC saline for 120 days, the sheet resistance of the silicone sheet, after DI water cleaning, is greater than 120 Mega ohms.

A lab made roller is used as the special tool to realize the multi thin film layer coating.

The roller coating can reduce the thickness of the coating film to 30 to 50 µm for each single layer. The rolling coating also avoids air bubble problem which is very common for traditional dip coating.

For the typical non-hermetic micro-packages, three film layers are used; the first layer is thin FP 4450 which used to protect the IC chip and wire bonds as well as to provide mechanical support to the device. Then a layer of Parylene C is deposited to serve as the water vapor barrier with the thickness of 15 µm. The outer layer is Dow Corning MDX

4210 medical grade silicone to make the implantable device biocompatible with interface tissues. The silicone and FP 4450 are all rolled onto the telemetry circuit using the roller, and the total thickness of each material is less than 100 µm. After cure of the silicone, the whole device is considered fully packaged. Table 7 list all the candidate materials and coated thickness used in multi material multi film layer packaging experiment.

Table 7. Summary of packaging materials and method

Material Thickness Applying Comments

Silicone ~150 µm Rolling MDX 4210 Dow Corning

Parylene ~15 µm Deposited Parylene C

Epoxy ~150 µm Rolling FP 4450 Henkel

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Figure 23. Experiment seetup and block diagram of the receiver circuit

The experiment setup for lifetime measurement is shown in Fig.23. The output signal is observed by spectrum analyzer and oscilloscope. The block diagram of the receiver circuit is also shown in the plot. The receiver has three major parts, receiver coil, RF signal amplifier and pulse shaping block. Fig.24 is the photos taken from the oscilloscope of the received RF signal after modulation. The Fig.24 (b) is the zoom in view of the low duty cycle pulse in Fig. 24-a.

The packaged devices are immersed in 40 oC or 85 oC saline baths fofor simulated life time evaluation. The initial group of devices without Parylene coating has a life time over 70 days in the 40 oC saline bath. The test results are plotted in Fig.25. The current consumption increased from 0.55 µA to 0.67 µA. The increase is about 22%. This is because the vapor penetrates through the packaging and

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b) Figure 24. Photos of received low duty cycle RF signal the surface insulation resistance. The voltage supply is constant, so the power consumption of the device increased about 200% after 70 days test. TABLE II summarized the output signal shape change of the #10 device. These parameter changes also indicate a power consumption increase of the device after 70 days in saline solution.

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Figure 25. The pulse duration and current consumption versus time of #10 telemetry unit After the experiment on initial group of devices, the second group of devices with the water barrier layer of Parylene C has been put into 40 oC saline for test. After 30 days test, the second group of device didn’t show any obvious change on the output signal, the life time of this group of device is expected to be llonger than two months.

t1 t2

Beginning of the test 404µs 517 ns

End of the test 367 µs 586 ns

Table 8. The pulse period, width at the beginning and end of the test

From the test results discussed above, the non-hermetic, multi film layer, multi material packaging is proven to be an effective and convenient method for short to long term implantable devices. The #10 device lasted more than two months without the Parylene water vapor barrier layer. Fig.26 is the block diagram of the telemetry unit with RF power charging and remote controlling function. The new IC passed functional test, the

103 performance of this circuit after packaged with all three materials mentioned above using multilayer coating method is being evaluated and will be reported when the results are analyzed [II-18].

Figure 26. Block diagram of the telemetry circuit with RF power and remote controlling

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Conclusion

In thesis part II, the theories of non-hermetic micropackage and thin film multilayer coating are proposed based on previous research. The single material package experiments, using FP 4450, MDX 4210 silicone and Parylene C, proved the feasibility of this non-hermetic package and also testified the expectation of the high performance of the proposed multi-layer coating approach. A previously designed micro-power telemetry unit is assembled and characterized for sensing pressure, potential, humidity, and resistance changes in the implant site. The non-hermetic micropackage techniques were used to package functional telemetry units and tested in 40 oC saline bath to have a live time longer than 60 days. The experiments demonstrated the usefulness of non-hermetic micropackage technology for packaging implantable biomedical microsystems for research and health care.

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Appendix Coating tools:

Lab made roller (3/8 inch ID silicone tubing, 3/8 inch resistor), small brush, and Teflon board, Teflon tip.

Coating procedure:

1. Preparation of PCB and wires

Each single piece of PCB should be well cut with the length of 2.5 cm and 1.5 cm in width. The finger gap is 0.5 mm. Copper thickness on PCB is 20 µm.

1) Use sand paper # 250 to round the four corners of the rectangular PCB, then

zero grade sand paper to smooth the edge of PCB.

2) Choose twisted pair wire with appropriate stiffness( to match the mechanical

impedance with PCB )

3) Strip both ends of the two wires, 1 inch ( to be solder onto PCB ) and 0.5 inch

( to be solder with connector) separately.

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4) Solder wires to PCB carefully. Make sure the soldering spots are flat and

smooth.

2. MDX

1) MDX and curing agent are mixed at 10:1 ratio (15-20 gram, 1.5-2.0 curing

agent), and then leave the mixture in vacuum for 10 minutes to remove air

bubble.

2) Pick some MDX place on the Teflon board, use roller roll the MDX on Teflon

board until the MDX uniformly distributed and the layer is thin enough (~ 100

µm). Roll the MDX onto the surface of PCB, use our roller, each roll will

accumulated a thickness of 50-100 µm, if push the roller hard (don’t deform

the roller) the thickness goes down to about 30 µm.

3) Half cure the coated MDX (40oC, 1 hr)

4) Repeat 2), 3) until reaching the desired thickness

5) Cure the MDX (40oC,2hrs)

3. ENCAP

I. Coated as soldering area protection

1) Apply some ENCAP according to numbers of PCB are going to be coated in a

small dish, and then degas the ENCAP in vacuum for 20 minutes.

2) Brush soldering area with ENCAP until the whole area is covered. Put the coated

PCB in vacuum for 20 minutes to remove air bubbles.

3) Partial-cure the ENCAP (80oC, 2 hrs).

4) Repeat step 2) and 3) two to four times until the soldering area is totally and

covered and the surface of the ENCAP is smooth.

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5) Cure (90oC, 5 hrs).

II. Coated as coating layer

1) Degas the ENCAP 5- 10 minutes (optional)

2) Pick some ENCAP place on the Teflon board, use roller roll the ENCAP on

Teflon board until the ENCAP uniformly distributed and the layer is thin enough

(~ 100 µm).Roll the ENCAP on the surface of PCBs, the thickness of each roll

depends on the pressure applied on the roller.

3) Half cure the coated ENCAP (80oC, 2 hrs).

4) Repeat 2), 3) until achieve the desired thickens.

5) Cure the ENCAP (90oC, 5 hrs).

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