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Electronics design and in vivo evaluation of a wirelessly rechargeable fetal micropacemaker
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Electronics design and in vivo evaluation of a wirelessly rechargeable fetal micropacemaker
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Content
ELECTRONICS DESIGN AND IN VIVO EVALUATION
OF A
WIRELESSLY RECHARGEABLE FETAL MICROPACEMAKER
by
Adriana Nicholson Vest
A Dissertation Presented to the
FACULTY OF THE USC GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfillment of the
Requirements for the Degree
DOCTOR OF PHILOSOPHY
(BIOMEDICAL ENGINEERING)
December 2015
Copyright 2015 Adriana N. Vest
ii
To my family.
1
Acknowledgements
I would like to extend my deepest appreciation to those people who have enabled
me to dedicate myself to this work and have helped me along the path professionally
and personally. This research would not have been possible without their support.
I thank my advisor, Dr. Gerald Loeb, for his encouragement and guidance that
enabled me to grow into the scientist and engineer that I am today. His creative ideas
and passion for imparting knowledge have inspired me to tackle difficult problems in
theory and experimentation. Through his high expectations and with patience, he
pushed me to develop professionally and mentally to levels that I was not aware I
could achieve. Dr. Loeb, thank you for investing so much of yourself in my success as
a researcher.
I would like to thank my committee, Dr. Ellis Meng, Dr. Terence Sanger, Dr. James
Weiland, and Dr. Yaniv Bar-Cohen, for their efforts in keeping me focused on the right
questions and doing the right work to lead me to my goal. In addition to their
mentorship, they have also led me by example and have served as templates for the
researcher I desire to become.
My labmate and friend, Li Zhou, has made this research possible in a very
significant way. His expertise on materials and mechanics has enabled the
micropacemaker to take form. He has worked diligently through fabrication and pre-
2
clinical testing, collecting data that I have analyzed within this body of work. Without
his companionship and partnership throughout this project, the obstacles for success
would have been insurmountable.
I also have had the privilege of working with the most collaborative group of
research clinicians I have ever met, including Dr. Ramen Chmait, Dr. Yaniv Bar-Cohen,
Dr. Michael Silka, Dr. Jay Pruetz, Dr. Erlinda Kirkman, and Dr. Catalina Guerra. Thank
you for your dedication to this work and your guidance to me throughout. Thank you
to Dr. Ann Mohrbacher, who has made this research possible in a very special way.
Thank you to my colleagues and collaborators within the Medical Device
Development Facility, SynTouch, and General Stim. I appreciate all of the input that
has been given to me throughout the project from these teams, including technical
advisement, development resources, and a multitude of hours of time and effort. To
Xuechen Huang, your fresh perspective on difficult theoretical concepts has given me
new energy to tackle those problems on multiple occasions. To Kaihui Zheng, I am
especially grateful for your introduction to the fetal pacemaker project and for the
chance to build upon the foundation built by you. I am also grateful to Gary Lin, Ray
Peck, Michael Lu, Viktoria Norekyan, Hithesh Reddivari, Shane Garcia, Jon Srdel,
Petcharat “May” Denprasert, and Sam Kohan, all of whom helped technically in many
experiments. Thank you to Dr. Francis Richmond, Shraddha More, Kunjan Shah, and
Anajana Krishnan for your expertise in regulatory science, which has helped to make
this technology more clinically feasible as a lifesaving tool. I am also grateful for many
conversations with colleagues Mandy Lai, Zhe Su, and Shanie Liyanagamage, who
3
inspire me with their own accomplishments and have provided me with a technical
sounding board to discuss new ideas and research directions.
I would like to express my appreciation for the entire community at USC, which
has provided me with a fertile environment in which to develop, as well as support
resources that have kept me able to dedicate my time to research. The community at
USC and in the Viterbi School of Engineering has exposed me to people from all over
the world with diverse backgrounds and experiences. It is here that I have had the
opportunity to immerse myself in an international culture without leaving the
country. I am also grateful for my graduate advisor Mischal Diasanta, who has
provided advice, encouragement, and friendship throughout the PhD process. The
project would also not be possible without the generous funding received from the
Southern California Clinical and Translational Science Institute (CTSI), the Robert E.
and May R. Wright Foundation, the National Institutes of Health (NIH), and the USC
Coulter Foundation.
My path to graduate studies was helped by many educators throughout my life.
Thank you to Dr. Robert Galloway, Dr. Thomas Withrow, Dr. Michael Goldfarb, Dr.
Brad Wood, Dr. Duco Jansen, and Dr. Anita Mahadevan-Jansen, who brought me into
their labs as a fledgling scientist over the summers and semesters during my
undergraduate years and made it possible to start building skills that enabled me to
jump into my PhD studies. I am also eternally grateful for those that shaped my very
earliest of experiences as a scientist, including Marty Howe, Helen Francis, Michael
4
Cerio, Virginia Blenke, Linda O'Berry, Angela Ivey, and Kathleen King, all of whom
inspired and fueled my interest in science and who made me believe I was capable.
Finally, on a more personal note, I want to thank all of my family and friends who
have supported me throughout the years. Thank you to my parents and my
grandparents, who have been my primary role models, providing me with my first
examples of hard work, dedication, selflessness, and generosity. It is through their
sacrifices and love that I was able to choose to dedicate myself to learning. To my
sister, Marisa, who keeps me grounded while encouraging me in any endeavor that I
should choose. To my parents-in-law and sisters-in-law, thank you for encouraging
me as one of your own.
Lastly, thank you to my husband, Dylan Vest, who has supported me throughout
this process. Thank you for being everything that I need at any given time with very
little notice. It is a big feat you accomplish every day. You celebrate my
accomplishments, push me forward when I do not think I can go any further, and
remind me to take a break when I have been pushing myself too much. I couldn’t have
done this without your love and encouragement.
5
Table of Contents
Acknowledgements ................................................................................................................................ 1
Table of Contents .................................................................................................................................... 5
List of Tables ............................................................................................................................................. 9
List of Figures ........................................................................................................................................ 10
Abstract ................................................................................................................................................... 13
Chapter 1: Introduction ..................................................................................................................... 16
Motivation ....................................................................................................................................... 16
Outline ............................................................................................................................................... 17
Chapter 2: Review of Applicable Technology in Cardiac Rhythm Management ......... 19
Introduction .................................................................................................................................... 19
Indications ....................................................................................................................................... 20
Heart Block............................................................................................................................... 20
Other Indications ................................................................................................................... 22
Biophysics and Physiology ........................................................................................................ 23
Pulse Polarity .......................................................................................................................... 24
Electrode-Tissue Interface ................................................................................................. 25
Strength-Duration Relationship ...................................................................................... 29
History .............................................................................................................................................. 31
Implantation Procedure ............................................................................................................. 40
State of the Art ............................................................................................................................... 41
Output Pulse Parameters ................................................................................................... 45
Pacemaker Code ..................................................................................................................... 46
CRT – Cardiac Resynchronization Therapy ................................................................. 49
Current Developments in Pacing and Leads ...................................................................... 51
Fetal Heart Block .......................................................................................................................... 54
Chapter 3: Design and Testing of a Percutaneously Implantable Fetal
Pacemaker ........................................................................................................................................... 56
Preface .............................................................................................................................................. 56
Abstract ............................................................................................................................................ 60
Introduction .................................................................................................................................... 61
Materials and Methods ............................................................................................................... 64
Mechanical Features ............................................................................................................. 64
6
Electronic Circuitry ............................................................................................................... 69
Electrode Properties ............................................................................................................ 72
Results ............................................................................................................................................... 75
In Vivo Percutaneous Implantation ................................................................................ 75
In Vivo Pacing Threshold .................................................................................................... 76
Power Optimization.............................................................................................................. 81
Discussion ........................................................................................................................................ 83
Implanting and Pacing a Fetal Heart .............................................................................. 83
Extending Functional Life .................................................................................................. 85
Device Longevity and Biocompatibility ........................................................................ 86
Clinical Translation ............................................................................................................... 87
Acknowledgments ........................................................................................................................ 89
Chapter 4: Electronics, Pacing Thresholds, and Power Budget ......................................... 90
Preface .............................................................................................................................................. 90
Abstract ............................................................................................................................................ 92
Introduction .................................................................................................................................... 93
Device Design ................................................................................................................................. 95
Methods ............................................................................................................................................ 96
Test Results ..................................................................................................................................... 98
Pacing Threshold Measurement ...................................................................................... 98
Power Budget ........................................................................................................................ 101
Discussion ...................................................................................................................................... 103
Acknowledgments ...................................................................................................................... 104
Chapter 5: Development and Testing of a Closed Loop Recharging System .............. 105
Preface ............................................................................................................................................ 105
Background ................................................................................................................................... 105
Investigation Into Component Values ................................................................................ 108
Ratio Between Bias Resistors ......................................................................................... 108
Value of RC and Minimum Current ............................................................................... 115
Accepted Component Values .......................................................................................... 116
Recharging During Fabrication: Benchtop System ........................................................ 117
Design....................................................................................................................................... 117
Testing ..................................................................................................................................... 118
Next Generation Design: High Power Driver System ................................................... 120
Chapter 6: Design and Testing of a Transcutaneous RF Recharging System ............. 121
Preface ............................................................................................................................................ 121
Abstract .......................................................................................................................................... 121
Introduction .................................................................................................................................. 122
Design .............................................................................................................................................. 127
Electromagnetic Field Generation ................................................................................ 127
7
Recharging Control Scheme ............................................................................................ 132
Calibration .............................................................................................................................. 135
Methods .......................................................................................................................................... 136
Results ............................................................................................................................................. 138
Magnetic Field and Current Generation ..................................................................... 138
Calibration .............................................................................................................................. 139
In Vivo Recharging .............................................................................................................. 142
Discussion ...................................................................................................................................... 143
Acknowledgments ...................................................................................................................... 145
Chapter 7: Preclinical Testing and Optimization of a Novel Fetal
Micropacemaker.............................................................................................................................. 146
Preface ............................................................................................................................................ 146
Glossary of Abbreviations ................................................................................................ 146
Abstract .......................................................................................................................................... 147
Introduction .................................................................................................................................. 148
Materials and Methods ............................................................................................................. 150
Device Design and Function ............................................................................................ 150
Device Implantation ........................................................................................................... 151
Fetal Sheep Follow-Up and Device Recharging ....................................................... 154
Evaluation of Pacing ........................................................................................................... 155
Device Explantation ............................................................................................................ 156
Results ............................................................................................................................................. 157
Evolution of Acute Implantation Technique ............................................................. 157
Follow-up and Recharging ............................................................................................... 160
Necropsy and Histology .................................................................................................... 161
Discussion ...................................................................................................................................... 165
Conclusions ................................................................................................................................... 170
Acknowledgements .................................................................................................................... 171
Clinical Perspectives .................................................................................................................. 171
Chapter 8: Assessing Cardiac Capture and Safety Factor ................................................... 173
Preface ............................................................................................................................................ 173
Abstract .......................................................................................................................................... 173
Introduction .................................................................................................................................. 174
Method ............................................................................................................................................ 181
Results ............................................................................................................................................. 182
Sheep 4 .................................................................................................................................... 182
Sheep 5 .................................................................................................................................... 184
Sheep 6 .................................................................................................................................... 188
Sheep 7 .................................................................................................................................... 191
Discussion ...................................................................................................................................... 193
Acknowledgments ...................................................................................................................... 197
8
Chapter 9: Conclusions and Future Directions ...................................................................... 198
References ............................................................................................................................................ 200
9
List of Tables
Table 2-1 Pacemaker Programming Code ................................................................................. 47
Table 2-2 Leadless Pacemakers .................................................................................................... 53
Table 3-1 Summary of Electrode Performance Based on Location................................. 81
Table 4-1 Power Budget Over Different Conditions ............................................................ 101
Table 5-1 Component Selection Table ....................................................................................... 116
Table 7-1 Summary of Implants ................................................................................................... 158
Table 8-1 Summary of Results ...................................................................................................... 193
10
List of Figures
Figure 2-1 The Electrode-Tissue Interface ................................................................................ 27
Figure 2-2 Strength-Duration Curve ............................................................................................ 30
Figure 2-3 First artificial pacemaker .......................................................................................... 32
Figure 2-4 Early Pacemaker Schematics .................................................................................... 34
Figure 2-5 Demand Pacemaker Schematic ................................................................................ 36
Figure 2-6 Dual Chamber Pacemaker Schematic .................................................................... 37
Figure 2-7 The Evolution of Pacemaker Technology ............................................................. 51
Figure 3-1 Generation II Populated Pacemaker PCB ............................................................. 58
Figure 3-2 Generation III Pacemaker PCB ................................................................................. 58
Figure 3-3 Pacemaker Schematic .................................................................................................. 60
Figure 3-4 First Generation Fetal Pacemaker ........................................................................... 64
Figure 3-5 CAD Model of the Fetal Pacemaker ......................................................................... 65
Figure 3-6 Injection Molding System Schematic ..................................................................... 67
Figure 3-7 Pacemaker Circuitry Schematic ............................................................................... 69
Figure 3-8 Pacing Rate Over Time ................................................................................................ 71
Figure 3-9 Cyclic Voltammetry ....................................................................................................... 74
Figure 3-10 Percutaneous implantation site at post mortem exploration .................... 76
Figure 3-11 Strength-Duration Curves for Various Implantation Sites.......................... 78
Figure 4-1 CAD Model ........................................................................................................................ 94
Figure 4-2 Pacemaker Schematic .................................................................................................. 99
Figure 4-3 Strength-Duration and Charge ............................................................................... 100
11
Figure 4-4 Pacing Rate Over Battery Discharge Time ......................................................... 100
Figure 5-1 Relationship Between Output Pulse Rate and Supply Voltage .................. 107
Figure 5-2 Programmable Unijunction Transistor ............................................................... 110
Figure 5-3 Voltage Vs. Current in PUT ....................................................................................... 110
Figure 5-4 Relaxation Oscillator Operation ............................................................................. 112
Figure 5-5 Pacing Rate Vs. Supply Voltage for Different Ratios ...................................... 113
Figure 5-6 Discharge of 3 mAh Battery for Different Ratios ............................................. 114
Figure 5-7 Recharging Algorithm ................................................................................................ 119
Figure 5-8 Benchtop Recharging System Front Panel ......................................................... 120
Figure 6-1 Pacemaker Insertion Strategy ................................................................................ 123
Figure 6-2 Schematic and SPICE Simulation of Pacemaker .............................................. 124
Figure 6-3 Calibration Curve ......................................................................................................... 126
Figure 6-4 EMF Vs. Distance from Primary Coil .................................................................... 130
Figure 6-5 Schematic of the Coil Driver and Antenna ......................................................... 131
Figure 6-6 Diagram of the Recharging Strategy .................................................................... 132
Figure 6-7 Field Measurements In Vitro ................................................................................... 139
Figure 6-8 Effects on Calibration ................................................................................................. 140
Figure 6-9 Corrected Calibration Curves .................................................................................. 142
Figure 7-1 Fetal Implantation Equipment ............................................................................... 149
Figure 7-2 Ultrasound: Micropacemaker deployment ........................................................ 152
Figure 7-3 Ultrasound: Creating a Pericardial Effusion ...................................................... 153
Figure 7-4 Fetal ECG Showing Capture ..................................................................................... 156
Figure 7-5 Necropsy: Pacemaker Crossing Diaphragm ...................................................... 162
12
Figure 7-6 Histology At Implantation Site ............................................................................... 163
Figure 7-7 Radiograph of Micropacemaker ........................................................................... 164
Figure 8-1 Theory of Minimal Interval ...................................................................................... 179
Figure 8-2 Minimal Interval for Sheep 4 ................................................................................... 183
Figure 8-3 Minimal Interval for Sheep 5 ................................................................................... 186
Figure 8-4 MicroCT on Sheep 5 .................................................................................................... 186
Figure 8-5 Minimal Interval for Sheep 6 ................................................................................... 189
Figure 8-6 Minimal Interval for Sheep 7 ................................................................................... 190
13
Abstract
A fetal pacemaker can dramatically improve the outcome for fetuses that develop
complete heart block in utero. We have developed a rechargeable fetal
micropacemaker in order to treat severe fetal bradycardia with comorbid hydrops
fetalis. Implanting our fetal micropacemaker could reverse the typically fatal outcome
of this condition, resulting in the resolution of hydrops within one to two weeks by
pacing the heart and restoring adequate blood flow to the fetus.
The main requirements of the device are that it be implanted with a minimally
invasive technique and be implanted entirely within the fetal chest. The pacemaker
developed meets these requirements by being designed to fit within a standard fetal
surgical cannula, putting a hard constraint on the size of the device. The size limitation
dictates the volume available for circuitry and a power source inside the implant. A
simple fixed-rate and fixed-amplitude relaxation oscillator based on a single
transistor provides stimuli while also meeting stringent requirements for low power
consumption and low development cost. A commercially available cylindrical,
rechargeable 3 mAh lithium ion cell provides power.
In this dissertation, the limits of the simple circuitry and power source are
identified and compensated for. A power budget provides an analysis of the battery
14
life with any given combination of components. The main draw of current from the
battery is the output pulse, so it is desirable to set the stimulus strength as low as
possible to conserve power in order to maximize the recharge interval, but it is also
important to include a safety factor to ensure effective ventricular capture for
somewhat unpredictable electrode placements and tissue conditions. The unknown
conditions of each electrode placement leads to a need to monitor the electrode-
myocardium interface in order to determine that adequate pacemaker output is being
provided. This is typically accomplished by observing the minimal stimulus strength
that achieves threshold for pacing capture. The output of the micropacemaker cannot
be programmatically altered to determine this minimal capture threshold, but a
safety factor can be inferred by determining the refractory period for ventricular
capture at a given stimulus strength. This is done by measuring the minimal timing
between naturally occurring QRS complexes and successful stimuli. Upon pilot testing
this method in four fetal sheep, data demonstrate that a relative measure of threshold
is obtainable, providing valuable real-time information about the electrode-tissue
interface.
Limited volume inside the implant also constrains the engineering of the
recharging system, which relies on inductive coupling to provide wireless current to
the lithium ion cell. To overcome the lack of regulation circuitry within the implant, a
method for controlling the recharging process was developed and utilizes pacing rate
as a measure of battery state, a feature of the relaxation oscillator used to generate
15
stimuli. The verification of the recharging system shows successful generation of
recharging current in a fetal lamb model.
16
Chapter 1: Introduction
Adriana N. Vest
Motivation
The design, development, and testing of a medical device entails the application
of many knowledge domains, including electrical engineering, mechanical
engineering, materials science, regulatory science, and physiology. The body of work
presented here focuses on the electrical and physiological aspects in particular, but
considers also concurrent research and development on other aspects. Specifically,
this work details the investigation into requirements of the fetal pacemaker and the
many design iterations of the prototype to meet those requirements within the
constrained resources available in a university lab with a small engineering team.
17
Outline
This thesis describes the research done to accomplish the clinical application of
the fetal micropacemaker through the following chapters:
Chapter 2: Review of Applicable Technology in Cardiac Rhythm
Management
A review of the literature is provided on the clinical indications for
pacing, the electrophysiology of cardiac stimulation, a history of the
development of pacemakers and the future directions of pacemaker
technology.
Chapter 3: Design and Testing of a Percutaneously Implantable Fetal
Pacemaker
This chapter provides the basic need for a fetal pacemaker and the
general engineering considerations necessary to bring the device to
realization. Preliminary data is also presented to set the stage for future
experimental direction.
Chapter 4: Electronics, Pacing Thresholds, and Power Budget
This chapter delves deeper into the electronic design of the fetal
pacemaker, including an analysis of the power budget, a chief concern
given the very small lithium ion cell used as a power source.
18
Chapter 5: Development and Testing of a Closed Loop Recharging System
The recharging strategy is laid out in this chapter, and a low power
system is designed to evaluate the state of charge of the power source
wirelessly on the benchtop. An algorithm is implemented to conduct
recharging of the pacemaker automatically.
Chapter 6: Design and Testing of a Transcutaneous RF Recharging
System
A high power version of the recharging system is developed to achieve
transcutaneous recharging and testing is performed to implement the
system.
Chapter 7: Preclinical Testing and Optimization of a Novel Fetal
Micropacemaker
This chapter details the preclinical study of the pacemaker in a fetal
sheep model.
Chapter 8: Assessing Cardiac Capture and Safety Factor
A method is developed to analyze the quality of the electrode-tissue
interface and therefore the safety factor of the stimulus in vivo.
Chapter 9: Conclusions and Future Directions
This chapter summarizes the results from this body of work and
proposes expansion of the technology into additional indications. It
also discusses additional technological developments that could be
undertaken.
19
Chapter 2: Review of Applicable Technology in Cardiac Rhythm
Management
Adriana N. Vest
Introduction
Cardiac pacing is one of the most successful applications of implantable medical
devices. Its evolution owes much of its success to the rise of integrated electronics,
the ability of physicians to assess and apply technology, and the expanding patient
population that benefits from pacemaker use. The original indication for pacemakers
was symptomatic bradycardia, and permanent cardiac pacing remains the only
effective treatment for this chronic condition [1].
Commercial pacemakers include an analog sensing circuit, analog output circuit,
an application specific integrated circuit (ASIC), memory, bidirectional telemetry, and
a power source. However, in order to achieve effective pacing of heart rate, the only
necessary components are an oscillator, a power source and an output electrode.
Therefore, the first pacing devices were constructed from a blocking oscillator or
astable multivibrator oscillator, a battery, and a simple output electrode such as
stainless steel suture wire or a long needle [2].
20
In this thesis, I will review the indications for pacing, the implantation procedure,
the biophysics and physiology of electrical stimulation in the heart, the history of the
development of pacing, and the current state of the art in pacemakers. These topics
will give the necessary background for my thesis, which is to develop a miniaturized
and percutaneously implantable pacemaker for treatment of the original indication
of pacing but in a new and unusual type of patient: a fetus in utero with complete
congenital heart block and comorbid hydrops fetalis.
Indications
Heart Block
Pacemakers were originally designed for managing Stokes-Adams attacks or
syncope, which are symptoms of complete heart block and severe bradycardia. The
physicians developing pacemakers hypothesized that taking over the role of the
pacemaker cells in the heart would keep the heart beating at a physiological rate and
stop the attacks altogether [3, 4]. It became clear that pacing could treat the
underlying condition of bradycardia, and therefore other severe symptoms
associated with long term bradycardia, such as congestive heart failure, could be
alleviated by the device.
21
Heart block is not always severe enough to cause syncope, and can be classified in
three different degrees. This classification can be done with echocardiography or
electrocardiography (ECG). Evidence of first degree block includes a long PR interval
with no missed ventricular beats. First degree block may not present any symptoms
and does not always need treatment. Second degree block is evinced by an occasional
to semi-frequent failure of atrial depolarization followed by ventricular contraction.
Third degree block, or complete heart block, is diagnosed when there is no association
between the rates of atrial and ventricular contraction. This can be seen on an ECG as
QRS complexes that occur at a rate that is independent of P waves and usually at a
much slower rate.
The type of heart block also depends on where the conduction breaks down.
Proximal heart block occurs within the AV node and is usually benign because
pacemaker cells in the AV node just beyond the site of the block take over the rhythm
of the heart. When heart block occurs in the Bundle of His or below, called distal heart
block, the patient is in a more dangerous condition because the ventricular escape
rhythm tends to be unphysiologically slow and less reliable, often worsening over
time.
Heart block at any degree can develop in utero or at any point throughout life. It
can be caused by various pathologies and mechanisms, whose incidence varies with
age. Fetal heart block, or congenital heart block, is usually caused by either an
22
autoimmune disease of the mother or by structural anomaly in the developing heart.
An ECG can be difficult and unreliable to obtain for a fetus, and so the dissociation of
atrial and ventricular contractions is usually diagnosed with ultrasound. Heart block
presenting itself for the first time in an adult is usually the result of ischemic heart
disease.
The treatment of choice for symptomatic bradycardia and complete heart block is
implantation of a pacemaker that can speed up the contraction rate of the ventricles
allowing for adequate circulation of blood throughout the body[1]. Current
pacemaker technology and science provide adults and pediatric patients with an
excellent prognosis and a high quality of life post-implantation.
Other Indications
In addition to treating heart block, cardiac pacing is used today to treat Sick Sinus
Syndrome (SSS), the onset of tachyarrhythmias, heart failure, neurocardiogenic
syncope, and hypertrophic obstructive cardiomyopathy. Pacing has been shown to
terminate supraventricular and ventricular arrhythmias. Future indications may
include prevention of atrial fibrillation by atrial pacing, but this treatment is still
under investigation[1]. Sick sinus syndrome, also known as sinus-node dysfunction,
is actually the most common indication for permanent pacemaker implantation
today[1, 5]. It refers to a number of disorders of sinus rhythm that result from atrial
23
disease, including sinus bradycardia, sinus arrest, sinoatrial block, and paroxysmal
tachycardias (the bradycardia–tachycardia syndrome) [4-6].
Biophysics and Physiology
The biophysics of cardiac electrical stimulation are the same as any neural
prosthetic device. Electrical current is injected into the extracellular fluids
surrounding the excitable cardiomyocytes and creates a temporal and spatial voltage
gradient in the tissue. The gradient induces charge to flow across the cell membrane
by capacitive conductance in response to the quickly changing voltage gradient
created by the injected current. This changes the potential across the cell membranes
of the tightly coupled cardiomyocytes. Some of the cardiomyocytes will be
depolarized to the threshold for voltage-dependent ionic channels to sustain an
action potential. If a sufficient number of cardiomyocytes generate an action
potential, it will spread via the tight coupling throughout the heart, inducing a
functional contraction.
A myocyte has a normal cycle of depolarization and repolarization. In its resting
state, before depolarization, the cell membrane is a charged spherical capacitor. The
potential of the cell is about -70mV when measured from the inside with respect to
the outside. In order to elicit a depolarization and consequent contraction, the
membrane potential must be depolarized, or raised, by 15 to 20mV [7, 8]. This
24
depolarization results in a sequence of openings and closings of sodium, calcium and
potassium channels in the membrane, which results in flow of the action current. The
ionic current is conducted resistively to adjacent myocytes through gap junctions.
This allows the muscle to contract in a coordinated manner across the millions of
myocytes that make up the heart.
Pulse Polarity
An important consideration for electrical stimulation is the polarity of the pulse.
Cathodal stimulation is considered superior [9] because less energy is needed to
bring the cell to depolarize when the electrode creates a region of high negative
potential outside of the cell membrane. The membrane becomes depolarized at the
site of the stimulation and hyperpolarized in regions surrounding the depolarization
because the capacitive stimulation current that enters a cell in one region along the
cell membrane is balanced by another hyperpolarizing current in other regions [7].
Anodal stimulation, even though it is less efficient, does generate excitation at the
virtual cathodes that are formed adjacent to the region hyperpolarized by the locally
positive potential at an anodal electrode [10].
Unipolar pacing was used in the past with a cathode tip electrode and the metal
case of the pacemaker acting as an anode. Today, most pacemaker leads are bipolar,
which means they have a cathode tip and an anode ring electrode located nearby.
25
These two electrodes are generally about 1 to 3 cm apart, which is considered a large
distance in the field of electrical stimulation because of the relatively large volume of
tissue between the contacts [8]. The shape of the electric field generated at the
cathode is not significantly affected by the anode when the contacts are so spaced. As
predicted, clinical studies show minimal differences between unipolar and bipolar
stimulation [1]. However, bipolar has become the standard of care because bipolar
electrodes are better for sensing bioelectrical signals. An additional benefit of the
stimulation anode being on the lead instead of the case is to restrict the volume of
tissue through which the stimulus current flows and generates stimulus artifacts that
interfere with sensing. The bipolar configuration picks up near-field events with a
much better signal-to-noise ratio, and far-field events with much lower amplitude.
Bipolar configurations also have drawbacks though, including an increased physical
bulkiness and more opportunities for failure [1], so modern pacemakers have the
ability to change their programming to switch from use of bipolar electrodes to
unipolar electrodes [8].
Electrode-Tissue Interface
When a metal electrode is in contact with the extracellular fluid, an aqueous
electrolyte, stimulation current flow across the interface must transfer between the
two phases. Charge is carried by electrons in the solid metal phase and ions in the
aqueous phase. The interface allows the transfer of charge by transduction of energy
26
and uses two mechanisms, non-Faradaic reactions and Faradaic reactions. Faradaic
reactions involve the transfer of electrons across the phase barrier, causing non-
reversible reduction and oxidation reactions in the electrolyte and metal electrode.
During non-Faradaic reactions, charge is transferred through a reversible
redistribution of charged species in the space around the electrode, including the
migration of ions in the electrolyte to and from the electrode, adsorbed anions onto
the electrode, and polar molecules such as water reorienting to separate charge. The
redistribution causes a double layer of opposing charges to form without transfer of
electrons across the interface. These reactions are depicted in Figure 2-1. Faradaic
reactions can be represented by a purely resistive impedance which dissipates charge
(R Faradaic in Figure 2-1), whereas non-Faradaic reactions, also known as capacitive
charge transfer, can be modeled as a capacitor (C Double Layer).
27
Figure 2-1 The Electrode-Tissue Interface
The electrode tissue interface has two main mechanisms for charge transfer, Faradaic and non-
Faradaic, also known as capacitive charge injection [11].
Both the Faradaic and non-Faradaic reactions occur in parallel on a working
electrode, but their relative contributions depend on the parameters of the
stimulation pulses applied to them as well as the size and nature of the metal-
electrolyte interfaces. The non-Faradaic processes act like a capacitor, accumulating
voltage while charge flows in one direction and discharging while charge flows in the
reverse direction. Faradaic reactions occur very slowly until a threshold potential is
reached across the metal-electrolyte interface. The thresholds, or charge storage
capacities (CSC), are documented for specific electrode materials. For example,
platinum has a reversible charge storage capacity of about 50-300 µC/cm
2
. Activated
iridium has a CSC of 1-3 mC/cm
2
before it reaches charge densities and voltages that
28
might result in electrolysis or corrosion. Irreversible Faradaic reactions are
unfavorable because they change the chemical environment and can produce
byproducts that are toxic to tissue. In a properly designed biological stimulator, the
amount of charge that can flow in one direction before it is completely reversed is
carefully designed to avoid reaching the threshold for the Faradaic reactions. This is
done by the use of charge-balanced biphasic or capacitively coupled monophasic
pulses [10-12]. A model of the electrode tissue interface is presented in Figure 2-1
and includes a pair of Zener diodes with Zener voltages around 0.6V to 0.8V [13] to
represent the threshold voltage of an iridium oxide electrode. This pair of elements
models the strong nonlinearity in electrode impedance that occurs when potentials
exceed the threshold voltage of the electrode.
Activated iridium oxide is a particularly useful material for stimulation because of
its ability to convert between Ir
3+
and Ir
4+
states within the conductive oxide layer to
accommodate additional charge storage capacity [14]. In the pacemaker presented in
this thesis, the cathodal electrode is constructed by using pure iridium drawn wire
coiled into a corkscrew shape. The exposed tip is activated with cyclic voltammetry
to produce a multilayered oxide film. The oxide layer has the added benefit of
dramatically increasing the effective surface area of the electrode, which reduces
electrode impedance and improves energy efficiency and battery life of the stimulus
generator. For a monopolar electrode, electrical impedance tends to be reciprocally
related to the surface area of the electrode and the cross-sectional area of the tissue
29
through which the stimulus current must flow. A smaller electrode will have a higher
impedance and therefore will permit less current flow for a given stimulus voltage,
but the resulting current will be concentrated in a smaller volume of myocardium, so
it may be more effective in the case of cardiac pacing.
Strength-Duration Relationship
The relationship between the strength and the duration of the applied stimulation
pulse at threshold for tissue excitation provides important insights into the
underlying physiology. This relationship states that the stimulus amplitude necessary
to elicit an action potential in an excitable tissue increases as the duration of the
stimulus is decreased. Two points along this important relationship have been used
historically to characterize the target tissue. Firstly, the rheobase is the amount of
current amplitude necessary to excite the tissue with a pulse of infinite duration. The
second point is the chronaxie, which is the minimal pulse duration that achieves
excitation when the stimulus current is twice the rheobase value. Stimulus pulses
whose duration is in the range of chronaxie tend to be most efficient to deliver, in
terms of the total charge necessary to achieve excitation (the product of stimulus
current times duration for a simple square-wave pulse) and the voltage required to
deliver that charge [8, 10]. These durations reflect the membrane time-constant of
the cells being stimulated (chronaxie is 0.7 times the membrane time-constant [15]),
so the transmembrane currents are effectively integrated by the membrane
30
capacitance. Typical stimulus pulses to elicit a contraction of the heart deliver a
charge in the range of .1 to 50 µC over a duration of 0.1 to 2ms [8].
The strength duration relationship is influenced by many factors including the
electrode tissue interface, relative position of the electrode, epinephrine levels in the
body, tissue state (healthy, injured, scarred), etc. The relationship curve shifts due to
these factors and a higher threshold may be needed to excite the tissue at some times.
Therefore, a safety margin is incorporated into the setting of the pacemaker output
at the time of implantation. Thresholds usually increase after implantation by a factor
of 2, but then stabilize. Generally a 100% safety margin is used [8].
Figure 2-2 Strength-Duration Curve
The Example Strength-Duration Curve provided here shows the resulting data taken when strength of
stimulus or duration of stimulus is varied, and the minimal stimulus necessary to capture the
endocardium is recorded [9].
31
History
The development of cardiac pacing is well documented and involved several
groups of researchers, physicians and engineers working together to bring to fruition
a reliable medical device. Physicians led the pacemaker technology development in
close concert with engineers at the beginning in the 1950s and 60s. As technology
became more complex, manufactuers took over as the dominant influence in the
1970s and have been the main driver of advancement ever since [4].
The possibility of cardiac electrostimulation first became known near the end of
the 1800s. The first review article about electrostimulation of the heart was
published in 1889. The article reviewed a successful temporary pacemaker and a
successful use of cardiac defibrillation [6].
The first experimental pacemaker consisting of a mechanically operated electrical
source for impulse formation and rate was developed by Hyman (Figure 2-3) to treat
bradycardia and published in 1932 [6, 9, 11]. It delivered the electrical pulses to the
heart via a percutaneous needle to the right atrium.
32
Figure 2-3 First artificial pacemaker
This pacemaker developed by Hyman was powered by a crank (a.). A flow diagram shows the
operation of the system (b.) [11].
In 1952, Zoll showed that in an emergency situation, the heart could be captured
by using a plate on the chest and an esophageal electrode [6]. This was the first use of
electrical stimulation to treat Stokes Adams attacks. However, this transcutaneous
stimulation was intolerable by patients because of the high energy required to
stimulate the heart tissues through all of the adjacent tissues [9].
Very close to the same time Furman and Robinson were stimulating the heart to
contract endocardially with a unipolar electrode at the tip of a catheter. This
ultimately led to the endocardial stimulating electrodes we have today. The first
transvenous bipolar catheter was used for stimulation in 1958 by Furman [6].
33
The major breakthrough came when physician Åke Senning and engineer Rune
Elmqvist developed the first implantable pacemaker in 1958 [2, 4, 11]. This device
was well tolerated by the patient and avoided the danger of infection along the
percutaneous leads. The pacemaker delivered asynchronous ventricular pacing,
which was the first type of pacing used to treat symptomatic bradycardia. This early
pacemaker delivered 2 volt pulses over 1.5ms at a rate of 70 to 80 pulses per minute
[2]. Pacemaker output in early devices was usually several times higher than
necessary. The emphasis of the early devices was to drive the heart reliably. For
power, it used two 60 mAh Nickel cadmium rechargeable cells. It featured a coil
connected to the cells with a rectification diode for inductive recharging through the
skin. Elmqvist’s design used a vacuum tube RF generator at 150 kHz and a large
flexible primary coil, which was taped to the skin of the patient during recharging.
Inductive recharging was necessary about once a month and required an overnight
visit [2]. The stimulator aspect of the device used stainless steel suture wire with
polyethylene insulation as the electrodes. The electrodes eventually failed and were
replaced by more flexible models. Elmqvist eventually switched to mercury zinc
primary cells for power, since that was what most of the industry was using at the
time.
Elmqvist built the blocking oscillator device out of silicon transistors, which were
new to the market, and embedded the electronics in epoxy resin [2, 4, 9, 16]. It was
implanted in patient Arne Larsson, who went on to have nearly 30 pacemakers
34
implanted over his long lifetime, and who ultimately outlived Dr. Senning and Dr.
Elmqvist [17]. Another oscillator used by engineers in early pacemakers was the
complementary multivibrator with pnp and npn transistors. The mean drain for both
these systems was between 20 µA to 60 µA [9].
Figure 2-4 Early Pacemaker Schematics
(Left) Schematic of a blocking oscillator of original pacemakers. If the capacitor (C1) is charged, as
shown, the voltage on the base of V1 is positive with respect to the emitter, and no current flows
through V1. A small current discharges C1 through R1 and R2, and eventually reverses the polarity of
the condenser and turns the transistor V1 on. When the transistor starts to conduct, sudden current
flows in the primary, inducing current in the secondary and causing positive feedback because the base
becomes more negative. The collector current eventually plateaus (and the iron core of the
transformer may become saturated) at which point the secondary voltage disappears. This causes the
base to suddenly go positive with respect to the emitter again and transistor current is turned off with
a new cycle starting. C1, R1, R2, and the battery voltage determine the pulse frequency. The length of
the pulse depends on the transformer and emitter resistance. The pulses are fed to the electrodes over
transistor V2 with the DC component being blocked by C2 [18]. (Right) A complementary multivibrator,
astable [9] provides a second option for pulse generation.
35
Early pacemakers were fixed rate and delivered a pacing stimulus even if the
ventricles were generating spontaneous contractions some of the time. These
pacemakers were therefore characterized as asynchronous and competitive with
natural heart activity. Pacing during certain times in the contraction cycle induces
ventricular arrhythmias in some patients, so this competitive pacing was an
unfavorable aspect of early pacing [11]. This knowledge led to the development of the
demand pacemaker, which incorporates the use of a sense amplifier to disable pacing
when intrinsic activity of the heart is detected. An unintentional advantage of the
demand pacemaker is a lower amount of current draw from the battery because
fewer pacing pulses are delivered[11]. In 1963 Nathan et al developed a pacemaker
that was able to sense atrial beats and deliver ventricular beats after this sensed event.
This was the first application of an AV synchronous pacemaker [11]. Simply following
each atrial beat with a ventricular stimulation risks producing ventricular
tachycardia if there is atrial tachycardia, a common occurrence in patients with
complete heart block. This problem was overcome by including a refractory delay,
thereby establishing a maximal rate for ventricular stimulation. Nathan et al. also
programmed in a minimal ventricular pacing rate in case no atrial beats were
detected by the system [6, 19].
Barouh Berkovits, an electrical engineer active in the pacemaker field, introduced
another of the first versions of a demand pacemaker that was implanted in 1966[4].
The electrodes of the pacemaker both deliver the stimulus and sense an intrinsic beat.
36
The pulse generator that he developed included a sensing amplifier, a timing control
stage, and an output driver [11].
Figure 2-5 Demand Pacemaker Schematic
A demand pacemaker developed by Berkovits [11].
The first dual chambered pacemaker, which was introduced in the last part of
1970s, senses in one or both chambers and delivers pulses to both chambers.
Berkovits developed a pacemaker that paced both the atria and ventricles, but only
inhibited pacing if an intrinsic beat was sensed from the ventricles[11]. It was
described as an AV sequential pacer. Few units were implanted, however, because
this pacing modality was only appropriate for a small portion of the patient
population. It demanded more drain current from the battery and came in a bulkier
37
design than other pacemakers of the day. Berkovits’ design was a step on the road to
dual chamber pacing, dual chamber sensing, and dual inhibition and trigger (DDD).
Figure 2-6 Dual Chamber Pacemaker Schematic
A more advanced pacemaker by Berkovits that provides dual chamber pacing. [11].
In the late 1960s, most experienced implanters shifted to the transvenous
endocardial leads [4]. Transvenous leads encouraged a development in diagnosing
38
methods for arrhythmias, which in turn fostered the growth of the field of clinical
electrophysiology. It spurred new technologies to develop including implantable
defibrillators with transvenous leads and endocardial ablation of arrhythmia
generating conduction pathways [4]. This also led to a shift from surgeons to non-
surgical cardiologists performing the implantation of pacemakers.
In the 1970s, sick sinus syndrome became a new indication for pacing, lithium
battery technologies were embraced, and hybrid and integrated circuitry were
introduced. Lithium battery technology significantly outlasted the standard mercury
zinc cells and did not have their problems with self-discharge in quiescent circuits.
Lithium batteries also were much more energy dense and did not generate a gas
byproduct, allowing pacemakers to finally be hermetically sealed. Non-invasively
programmable pacemakers were also introduced in the 1970s, followed by
multiprogrammable units with bidirectional telemetry [4].
The first example of refined programming was released in 1972 by Cordis in their
Omnicor line of pacemakers. These used an integrated sensing amplifier and two
integrated digital logic circuits, and could be reprogrammed noninvasively for rate
and output. By using a handheld device over the skin where the pacemaker was
implanted, the physician could select a mode using a series of magnetic pulses that
vibrated a reed switch inside the pacemaker. A counter noted the number of changes
39
in position of the switch and associated the number with a corresponding value for
output [4].
In 1978 the Cyberlith was released as a direct competitor to the Omnicor line. The
Cyberlith was instrumented with even more modes of pacing, including sensitivity
settings. Its most novel development was a two-way telemetry system, which is what
is used today on modern pacemakers. The telemetry system was a product of the
collaboration between engineer Robert R. Brownlee and surgeon G Frank Tyers. This
bidirectional telemetry allowed the physician to adjust settings, but also provided
information about the stimulation rate, battery voltage, battery impedance, lead
impedance, and the integrity of encapsulation of the implanted device [4].
At the end of the 1970s and very beginning of the 80s, dual chamber pacing
devices were finally being developed and tested. The idea of restoring the
hemodynamic benefits of atrial contraction followed by ventricular contraction was
not a new one, but proved to be difficult to implement. The dual chamber pacemaker
allowed this development, but was difficult to program, hampering its acceptance by
physicians. It is now the standard of care for most patients receiving implants because
of its versatility and reprogrammability [4].
In 1981, a new generation of dual chamber pacemakers was released that was
able to sense and pace in both the atrium and ventricle. It paced the two chambers
40
sequentially when atrial rates were slow, but synchronously stimulated the ventricle
alone when atrial rates were present and adequate. The biggest impediment to
acceptance of dual chamber devices was implanting a second lead [4]. As leads were
reduced in size and implantation techniques continued to improve, this became less
of an issue. However a publication in 1984 by a review board of cardiac pacing
technology indicated that the implantation rate for dual chamber pacemakers was
significantly below the rate of new cases of cardiac diseases that could be treated with
dual chamber devices [4], consistent with slow clinical acceptance of the dual
chamber device.
Implantation Procedure
The implantation procedure for most patients consists of accessing the myocytes
of the heart with a lead that is threaded through the subclavian vein and into the right
atrium and right ventricle. This type of lead is classified as endocardial, and most
implants today are done with “active fixation” leads, which use a corkscrew
mechanism to anchor the lead into the myocardial wall. The operation begins with an
initial incision to create a pocket for the pacemaker electronics. This incision is placed
above the left pectoral muscle below the clavicle, and also allows the surgeon to easily
access the subclavian vein. The pocket is created by separating the fatty tissue and
fascia above the muscle layer with blunt tools. The vein is accessed with a trocar and
cannula, a catheter is placed into the vein and the tools are retracted. The catheter
41
allows the surgeon to thread the lead into the vein, which is done under fluoroscopy
guidance. Once the lead is in the correct location, the corkscrew is advanced using a
small tool that comes with the lead. The most common implant is a dual chamber
pacing device and usually involves placing an endocardial lead into the right atrial
appendage and another into the right ventricular apex [1].
Modern leads have a lumen that accommodates a wire to make the lead stiffer for
implantation. The wire is withdrawn after the lead is effectively affixed to the
endocardium. The wire can be straight, or the surgeon can switch the wire out for one
with a curved end to facilitate steering the tip toward the target.
Another method the physician uses to ensure proper location of the electrode
within the heart is to verify that the stimulation pulse is capturing the heart muscle
and determine the threshold of the electrode tissue junction. If threshold is unusually
high, the surgeon may conclude that the lead is not located properly, and can then
withdraw the corkscrew electrode and move the lead to another location. A suitable
threshold can be confirmed before the lead is permanently anchored to the tissue.
State of the Art
The improvements of implantable pacemakers from 1958 to today yielded a
device that is reliable, multiprogrammable, long lasting, small, and easy to implant.
42
Today’s pacemaker leads are much more robust than the stainless steel suture wire
used for the first implants and feature low impedance activated iridium cathodal
electrodes, better insulation, bipolar contacts, active fixation, steroid eluting
technology, and reduced failure rates. Pacemakers are mostly implanted
endocardially, which can be an outpatient procedure. Today’s battery of choice is the
lithium iodide battery, which has a much higher energy density than the original
mercury zinc batteries. Lithium batteries also allow the implants to be hermetically
sealed. The rise of integrated circuits during the same time that the pacemaker
evolved has yielded smaller generators that draw less current and feature full
microprocessor control, including features such as rate hysteresis.
Modern pacemaker topologies are extremely sophisticated and include analog
sensing and output, coupled with digital processing, programming, memory, and a
telemetry system [11]. The processor allows for the recording of electrocardiograms
recorded from within the heart via both the atrial and ventricular endocardial leads.
These have amplitudes from .5 to 20 mV based on the source of the signal. Bipolar
leads have better signal-to-noise ratios, but unipolar leads can also record. The
measured physiological signal has slew rates from 0.1-4 V/s and spectral power,
especially of the QRS complex, is concentrated between 10 and 30 Hz. Lower
frequency signals and high frequency noise are filtered out. Noise generators include
electronics outside of the body and electromyogram (EMG) signals from nearby
43
muscles [11]. A modern pulse generator has an expected device lifetime of five to ten
years, depending on how it is programmed [1].
Integrated rate responsive sensors and programming have also been developed
that allow patients to lead more active lives. The sensors can measure respiration,
body motion (accelerometers and piezoelectric crystals), cardiac contractility (QT
sensor, impedance), changes in central body temperature, venous oxygen saturation,
or EMG activity [6, 11]. These measurements reflect physical activity level to that the
heart rate can be modulated accordingly. None of these sensors are ideal though,
because of a lack of specificity, sensitivity, or speed. Rate responsive pacing has been
shown to improve patients’ quality of life, but relies on skilled programming of the
rate response in order to achieve the best effect [1].
Pacemakers can also now determine their own lead impedance and threshold [6].
The Auto Capture function, available in some units, allows the pacemaker to evaluate
the threshold of the tissue and deliver a safety factor above that threshold without
using excess battery power [6]. In addition to the developments in pacemakers today,
implantable cardiac defibrillators (ICDs) are now acting as pacemakers as well,
especially because many patients who present indications for an ICD eventually
develop sinus node dysfunction, atrial tachyarrhythmia, or other indications for
pacing [1]. ICDs with pacing functionality can sometimes terminate ventricular
tachyarrhythmias without excessive defibrillation shocks by pacing the heart faster
44
than the tachycardia. Pacing is also used to prevent ventricular tachyarrhythmias, in
intractable hypertrophic obstructive cardiomyopathy, and to prevent and interrupt
atrial fibrillation when it is not a result of bradycardia. The most recent clinical
research has been to establish atrial pacing as a treatment to prevent atrial fibrillation
in patients without bradycardia, but this treatment remains controversial today.
Another development that was spurred directly from the complex sensors that
make up pacemakers today is the implantable loop recorder (ILR), a version of the
wearable Holter ECG monitor used to diagnose intermittent arrhythmias. These are
cardiac diagnostic tools that are implanted in patients with difficult to diagnose heart
conduction issues. They record ECG information for up to two years, and allow
physicians to better understand underlying conditions [6].
Accepted indications for implantable cardiac devices are being extended even
further. Pacemakers are now being used to increase cardiac output in heart failure
patients with dyskinesia by pacing bi-ventricularly, in both the right and left
ventricles. This type of pacing is called Cardiac Resynchronization Therapy (CRT)
because the ventricles are contracting together instead of the right-to-left pattern
induced by pacing the right ventricle alone. The heart is ultimately more efficient at
pumping when the muscle is contracting together. Direct pacing of the left ventricle
is achieved by placement of a lead into the coronary sinus.
45
Output Pulse Parameters
Pulse generators are designed to produce stimulation pulses that consistently
provide adequate energy to the tissue to elicit a contraction. The parameter that most
effects this stimulated excitation is the charge delivered by the pulse, which is the
integral of the current flow over time. A current regulated generator, which uses a
current source between the active electrode and the reference electrode and delivers
a square pulse of current for a pre-set duration, would provide charge by the
relationship:
Charge (Coulombs) = Current (Amperes) x Pulse Duration (Seconds)
A regulated voltage output is used more frequently in modern cardiac
applications [8], and holds the potential difference between the active electrode and
reference electrode during the pulse, resulting in the current fluctuating as a function
of the reactive and nonlinear electrode impedance. The constant voltage output
requires less current regulating circuitry and therefore dissipates less power by that
means, but can be very inefficient in cases where the output voltage is much smaller
than the supply voltage.
The most efficient method of regulating output stimuli is by constant charge,
achieved by delivering a set amount of charge to the tissue regardless of impedance
changes. The charge is may be delivered by discharging a capacitor, but could also
employ a charge-metering means [20]. The fetal micropacemaker in this thesis
46
utilizes this constant charge method of delivery by charging up an output capacitor
(Cout) through a resistor (RC) to a pre-determined voltage (gate voltage on the PUT
Figure 3-3). The amplitude of the charge built up on the output capacitor and injected
into the tissue is determined by the value of Cout and the potential drop experienced
by Cout when it is discharged through the tissue, equivalent to the gate voltage minus
a small potential drop across the transistor (PUT). The shape and duration of the
actual stimulating current flow is essentially that of a capacitor discharging through
the tissue resistance. This is highly efficient as long as the time constant of this charge
delivery is less than one time constant of the target cells. In the case of cardiac
myocytes, this time constant is on the order of 0.5ms, as determined from chronaxie
extracted from the relationship between stimulation current required to achieve
capture for various duration pulses (see Figure 2-2).
Pacemaker Code
The ability to program devices and apply different methods of pacing has led to
the development of an international code to describe the form of pacing for a given
pacemaker. The code was started at the request of the Inter-Society Commission for
Heart Disease Resources by Drs S. Furman, N.P.D. Smyth and Parsonnet to refer to the
mode of pacing [4].
The code consists of five letters identifying the chamber that is paced, the chamber
that is sensed, whether a sensed event results in a trigger or a inhibition, the presence
47
of rate responsive pacing, and the chamber or chambers in which multisite pacing is
delivered [1]. The lettering system is shown in Table 2-1 below.
The most commonly used pacing modes are: AAI(R) single-chamber atrial pacing
without (or with) rate response, VVI(R) single-chamber ventricular pacing without
(or with) rate response, and DDD(R) dual-chamber pacing without (or with) rate
response [11]. Sick sinus syndrome (SSS), which disables the sinus node from
generating a pulse for the rest of the heart, would need atrial pacing, whereas AV
block, or heart block, would need ventricular pacing. As an example of the appropriate
pacing mode for a particular condition, SSS with AV block would need dual chamber
DDD pacing. Most physicians use two leads because the technology is available and
the reimbursement rate is substantially higher, so DDD pacing is the most common
mode today [1].
Table 2-1 Pacemaker Programming Code
The NASPE BPEG code for pacemakers describes the chamber paced, the chamber sensed and the
response to sensing [21] . The fetal micropacemaker would be a VOO pacing system.
48
Despite the numerous modes available, single chamber right atrial pacing (AAI)
might be adequate for the majority of pacemaker patients with sinus node
dysfunction and intact atrioventricular conduction. However, AV block develops in
0.6% to 5% of patients with SSS every year [1], so physicians opt for dual chamber
devices. For patients with only AV node problems, VVI pacing could be used instead
of dual chamber devices, but patients paced with VVI more frequently report
symptoms that suggest poor hemodynamics, including hypotension, pulsations in the
neck or abdomen, and headaches. . Dual chamber devices that are programmed
correctly assure maintenance of AV synchrony.
AV synchrony is important because it restores the heart to a physiologic state.
Loss of AV synchrony might reduce cardiac output by as much as 30% [1]. Physiologic
pacing is the most beneficial to those with intact AV conduction. It involves sensing
what electrical activity is still viable in the heart and using it to establish a good
synchrony between the atria and ventricles. In multiple randomized studies, there
was a decreased frequency of the complication of atrial fibrillation with atrial pacing
in patients with sinus node dysfunction, but these studies do not definitively show a
reduction in stroke, heart failure, and mortality with atrial pacing, contrary to
expectation [1]. In any case, physicians prefer to restore function as physiologically
as possible.
49
Another option for programming is VDD pacing with a single lead, which would
mostly be useful for patients with AV block. There is concern amongst clinicians and
researchers though that atrial sensing with a VDD lead, which has a floating atrial
electrode, is not reliable and that the lead is too wide [1]. This is another reason why
dual chamber DDD pacing is the accepted standard of care.
Because atrial tachyarrhythmias are common among patients with implanted
pacemakers, it is important for an atrial driven ventricular pacing system to be able
to switch to non-triggered pacing when an atrial arrhythmia presents itself. Modern
DDD pacemakers incorporate mode switching. This is important because if atrial
tachycardias were tracked, they would result in ventricular pacing that is too frequent.
Advances in programming and sensing have allowed these DDD pacemakers to
determine when an atrial arrhythmia is being sensed and to automatically switch to
a more appropriate pacing method for the duration of the arrhythmia episode [6].
CRT – Cardiac Resynchronization Therapy
A newer indication for pacing is heart failure with a conduction abnormality. This
therapy is often called Cardiac Resynchronization Therapy (CRT) and is a topic of
intensive research at the moment. CRT involves pacing to alter the degree of atrial or
ventricular electromechanical asynchrony and is done by pacing more than one atrial
or ventricular site (biatrial or biventricular pacing) or by pacing in an atypical
50
location like the left ventricle. CRT is helpful because the conduction delays that may
occur in heart failure patients result in an inefficient left ventricular contraction,
shortened diastole, and worsened functional mitral regurgitation. CRT can improve
the sequence of contraction, improving ventricular output. Research shows that
ventricular CRT has had a significant effect on improving heart failure symptoms [1].
In order to pace the left ventricle without introducing hardware into the chamber,
which could result in dangerous clots in the systemic circulation, an additional lead is
threaded through the coronary sinus and into a lateral or posterolateral cardiac vein
on the epicardial surface [1]. To address the extra challenges of placing leads near the
left ventricle, lead manufacturers have engineered various catheters, sheaths, and
guidewires.
CRT has also been suggested for preventing atrial fibrillation, but this is more
controversial than its use in treating heart failure. New algorithms aim to reduce
delays between electrical conduction in order to prevent fibrillation from happening
[6, 22]. Acceptable indications for CRT issued by the New York Heart Association
include a broad QRS complex, some left ventricular dysfunction and dilation, with
class III or IV heart failure. Cardiomyopathy, which can result from long term
ventricular apical pacing, is the next possible indication for CRT [6].
51
Figure 2-7 The Evolution of Pacemaker Technology
Today’s developments in pacemakers by industry leaders focus on miniaturization. This illustration
from St. Jude Medical, an industry leader in cardiac rhythm management, shows the general evolution
of pacemakers. (Image provided courtesy of St. Jude Medical, Inc.)
Current Developments in Pacing and Leads
Today, researchers continue to work on improving rate-response algorithms,
functionality of implantable cardioverter defibrillators, combinations of sensors to
optimize physiological response, advances in lead placement and extraction [1], and
processing algorithms that look at the details of the electrogram [11]. Within the past
few years, major manufacturers have begun to discuss development plans for
“leadless pacemakers”. This is because the lead is still seen as the weak point in
current pacemaker technology. Complications associated with leads include
dislodgment, insulation failure, and conductor fracture. These lead related issues can
also result in high current drain from the battery because of increased stimulation
52
thresholds at the electrode tip or leakage of stimulation current along the lead, which
are the most common causes of premature battery depletion [1]. While other parts of
cardiac rhythm devices are easily removed in the case of infection or failure, the
adhesions formed on chronicly implanted leads complicates their removal [23].
Several recalls due to insulation failure in the last few years highlight the fact that
leads can still be improved [24].
Leadless, modular devices (endocardial and epicardial) such as the one developed
by Nanostim, which was recently aquired by St. Jude Medical, have the benefit of being
able to be implanted with less invasive surgery. Therefore, these devices expand the
types and numbers of patients that can use a pacemaker. With their small size, around
1 cubic centimeter in volume, these devices pose a challenege for achieving the
battery life, multichamber sensing and stimulating systems that have become the
norm for conventional pacemakers.
53
Company Nanostim
(St. Jude
Medical)
Medtronic EBR Systems
Photo
Stage in
Development
First in Man Pre-clinical
development
First in Man
Volume
1cc 1cc 0.28cc
Energy
8 Years (Li
Primary Cell)
100% Pacing
7 Years (Li
Technology)
Unpublished
(Ultrasound Energy
from a secondary
implant)
Communication
Method
Conductive RF
Ultrasound
Table 2-2 Leadless Pacemakers
Evaluation of leadless devices in development and testing today. Medtronic device image obtained
from [25].
Because of the nature of industrial research and development, few details are
available about current technologies. However, published patents can be searched to
glean information about the leadless devices from the various groups. The summary
of information about these devices is provided in Table 2-2.
Although multiple companies have discussed to some extent their technology,
most publically available information is about St. Jude Medical’s Nanostim device.
54
Assuming the use of the best available medical grade primary cell energy densities
(lithium iodide; 0.9Wh/cc [26]), and a typical pacemaker drain current of 10uA when
pacing continuously [27], the Nanostim would be able to function for about 3 years.
This contrasts with the publications that claim an 8 year device lifetime. To achieve
this lifetime, the Nanostim could be using low-power, custom integrated and highly
efficient circuitry along with simple control schemes, similar to that shown by
Berkovits (Figure 2-5), with a VVI only capacity. Their estimation of battery life also
may not include any communication with the device or reprogramming.
Fetal Heart Block
Although pacemakers have addressed heart block in adults, this technology has
not been available to treat heart block before birth. Left untreated, heart block in the
fetus can progress into life-threatening congestive heart failure resulting in hydrops
fetalis [28-31]. Once hydrops fetalis develops, if the fetus cannot be delivered due to
prematurity or other clinical concerns, fetal demise is nearly inevitable[29].
Typical triage of fetal heart block consists of determining the cause of the block
and the degree. If the block is not caused by structural abnormality, pharmacological
and immunological therapies might be delivered systemically via the mother in order
to stop a hypothetical immune response from damaging the cardiac tissue of the fetus.
However, this regimen has not been significantly effective in clinical trials [32-34].
55
Therefore, there is no acceptable treatment for congenital heart block, and it can
progress from first or second degree to complete rather quickly and unpredictably.
Once complete heart block occurs, hydrops fetalis, the fetal equivalent of congestive
heart failure, follows in about 15% of cases [29].
It has been proposed that fetal heart block can be treated with pacing, similar to
the treatment of adult heart block. Restoration of a normal ventricular rate and
cardiac output should allow resolution of hydrops in 1-2 weeks and permit an
otherwise normal gestation. A conventional pacemaker would then be implanted
shortly after delivery. Over the last two decades, several investigators have attempted
to place pacemakers in a fetus [35-37]. To date, however, there have been no
survivors of fetal pacing. Previous approaches have relied on the placement of a
pacing wire on the fetal heart with an extra-uterine pulse generator implanted in the
mother. This has inevitably failed because of lead dislodgement due to fetal
movement.
The following chapters describe a miniaturized pacemaker that has been designed
specifically to treat heart block in fetuses. It is small enough to be implanted entirely
in the fetal chest using minimally invasive techniques, eliminating the need for leads
that resulted in failure of previous attempts at fetal pacing.
56
Chapter 3: Design and Testing of a Percutaneously Implantable
Fetal Pacemaker
Gerald E. Loeb, Li Zhou, Kaihui Zheng, Adriana Nicholson, Raymond A. Peck, Anjana
Krishnan, Michael Silka, Jay Pruetz, Ramen Chmait, Yaniv Bar-Cohen
© 2013 Biomedical Engineering Society. Reprinted from:
Loeb GE, Zhou L, Zheng K, Nicholson A, Peck RA, Krishnan A, Silka M, Pruetz J,
Chmait R, Bar-Cohen Y. Design and Testing of a Percutaneously Implantable
Fetal Pacemaker. Annals of Biomedical Engineering 2013; 41(1):17-27. 663.
Preface
This peer reviewed publication states the electrical, mechanical, and clinical
requirements for the fetal pacemaker and the design proposed to meet those
requirements. This preliminary design of the device utilizes the efficient oscillator
design of the original pacemakers and allows the stimulus generator with its
electronics and battery to occupy a much smaller volume than in present state of the
art pacemakers. Another design feature of the device is its ability to be wirelessly
recharged through the skin, a technology that was abandoned by the pacemaker
industry in the 1960s in favor of larger primary cells.
My contribution to this research is the design iteration of the pacemaker printed
circuit board (PCB). This also involved choosing components that met the
57
requirements of the design that would fit within the space allotted to the electronics
in a way that made them manufacturable. Additional contributions on the analysis of
in vivo pacing threshold and power optimization will be discussed in the following
section.
The space set aside for electronics was determined by the shelf cut out of the
ferrite, which is about 0.118 inches by 0.090 inches. This shelf was positioned so that
the electronics rest between the battery pin and the output connection of the
pacemaker, which is a platinum rod. A hole was bored into the ferrite to allow the
platinum rod to extend through it and reach the end of the device. In order to further
save space on the printed circuit board, the largest component, the output capacitor
of 1µF, was also extended into this hole and only needs one pad on the PCB. Figure
3-1 shows the platinum rod being soldered to the capacitor, which extends off of the
PCB.
The PCB design encompassed three iterations. The first version utilized an epoxy
substrate with copper and gold metallization. Initial tests of this board with the
proposed epoxy encapsulation method showed a significant amount of discoloration
of the populated PCB from water adsorption. To minimize water adsorption and also
provide the encapsulation epoxy with a bondable surface, the second iteration used
an alumina substrate with screen printed gold metallization. This PCB did not discolor
under similar soaking tests, but due to the poor resolution of the layout, as shown in
58
Figure 3-1, it was difficult to keep the solder from wicking into adjacent pads, an
especially troublesome issue for wirebond pads.
Figure 3-1 Generation II Populated Pacemaker PCB
This is the second version of the pacemaker electronics PCB populated with components before
wirebonding. The photo also shows the fixture (top) used during this stage of development that holds
the platinum rod and output capacitor in place during reflow soldering.
Figure 3-2 Generation III Pacemaker PCB
Solid works model of the third generation populated PCB, side view (left) and top view (right). The
diagram shows placement of wirebonds and a horseshoe shaped bracket (left edge of board) for
anchoring the battery to the PCB and soldering the ground pads to the battery case.
59
The third iteration of the PCB (Figure 3-2) featured a 0.005 inch thick alumina
substrate with metallization applied using photolithography techniques. The
metallization included 4 layers of metal: titanium tungsten, gold, nickel, and gold. This
combination of deposited metals enables both wirebonding and reflow soldering to
the PCB. Using photolithography enabled the design to have trace widths and spaces
as small as 0.001 inches. Solder wicking was not an issue for this version of the PCB.
Selection of electronics involved finding parts with very small form factors and
minimal packaging. The programmable unijunction transistor (PUT) selected was the
2N6027, one of the only PUTs on the market. It was obtained in a bare die version,
without the packaging found around most surface mountable components. The lack
of packaging does mean that light exposure can exert a photoelectric effect on its
silicon semiconductor, so care is taken to keep light away from the pacemaker
electronics when making critical measurements. The rectification diode (D1) selected
is the 1N5711 and also comes in a bare die package. Resistors were chosen in a 0201
package. The tuning capacitor was chosen to be ceramic, best for working at
frequencies around our recharging frequency of 6.78MHz. This ceramic capacitor also
comes in a 0201 package. The output capacitor was chosen to be medical grade, dry
tantalum because of its stable properties and minimal drift. This type of capacitor is
difficult to miniaturize, and along with its large value, 1uF, it could not be made
smaller than a 0402 package.
60
Figure 3-3 Pacemaker Schematic
Schematic diagram and component values for current version of the low stimulus charge (2.6µC)
pacemaker electronics.
Abstract
We are developing a cardiac pacemaker with a small, cylindrical shape that
permits percutaneous implantation into a fetus to treat complete heart block and
consequent hydrops fetalis, which can otherwise be fatal. The device uses off-the-
shelf components including a rechargeable lithium cell and a highly efficient
relaxation oscillator encapsulated in epoxy and glass. A corkscrew electrode made
from activated iridium can be screwed into the myocardium, followed by release of
the pacemaker and a short, flexible lead entirely within the chest of the fetus to avoid
dislodgement from fetal movement. Acute tests in adult rabbits demonstrated the
range of electrical parameters required for successful pacing and the feasibility of
61
successfully implanting the device percutaneously under ultrasonic imaging guidance.
The lithium cell can be recharged inductively as needed, as indicated by a small
decline in the pulsing rate.
Introduction
Complete heart block in the preterm fetus is a life-threatening emergency with no
effective treatment options beyond watchful waiting or preterm delivery [30, 31].
Fetal bradycardia due to heart block can progress in utero, and for more than a
quarter of these fetuses may result in hydrops fetalis [28, 38]. Once hydrops fetalis
develops, if the fetus cannot be delivered due to prematurity or other clinical
concerns, fetal demise is nearly inevitable [29].
Due to the often severe consequences of fetal heart block, various treatment
options have been undertaken in an effort to treat these fetuses. Pharmacologic
therapy has been trialed, usually with administration of medications to the mother
with anticipated placental passage of the agent to the fetus. Since damage to the fetal
cardiac conduction system is believed to be due to maternal auto-immune antibodies
in many of these cases [39], fluorinated steroids have been used to prevent or reverse
the heart block. Studies on this therapy, however, have demonstrated unclear, if any,
benefit from these agents [40]. Intravenous gammaglobulin has also been
administered to these patients in an attempt to reverse or halt progression of the
62
immunologic damage, but results of this therapy have also been disappointing [32,
33]. Beta-adrenergic agents are known to increase heart rates in children and adults,
but maternal administration of these agents results in heart rate increases that have
not been proven to affect overall survival [34].
When a newborn, child or adult presents with symptomatic complete heart block,
treatment usually consists of implantation of a pacemaker to ensure an adequate
heart rate. With appropriate pacing, these patients are usually asymptomatic with an
excellent prognosis. Similar benefits would be expected from pacing a fetus with
complete heart block, theoretically allowing resolution of hydrops in 3-4 weeks and
permitting an otherwise normal gestation. A conventional pacemaker would then be
implanted at delivery. Over the last two decades, several investigators have
attempted to place pacemakers in a fetus [35-37]. To date, however, there have been
no survivors of fetal pacing. Previous approaches have relied on the placement of a
pacing wire on the fetal heart with an extra-uterine pulse generator implanted in the
mother. This has inevitably failed because of lead dislodgement due to fetal
movement.
We have designed a single-chamber pacing system that is self-contained and can
be completely implanted in the fetus without exteriorized leads, thereby permitting
subsequent fetal movement without risk of dislodgement of the electrodes. Such a
design is now possible because of significant developments in medical device
63
miniaturization and advances in fetal surgical intervention, allowing the pacing
system to be percutaneously deployed through the maternal abdomen under
ultrasound and fetoscopic guidance.
The following requirements were identified by the clinicians and engineers as the
formal inputs to the design process:[41]
• Percutaneous implantation and anchoring of the electrode to the
pericardial surface of the myocardium
• Deployment of the pacemaker in the fetal thorax with a flexible lead to
accommodate cardiac and respiratory motion
• Use of a novel but simple packaging scheme to achieve required
longevity without bulky technologies
• Use of off-the-shelf electronic components to avoid expensive and risky
custom development
• Minimize voltage and current requirements to maximize functional life
without compromising pacing
• Ability to monitor charge-status of the implanted battery
• Inductive recharging system to extend functional life of the battery as
needed.
64
Materials and Methods
Mechanical Features
The fetal micropacemaker is designed for implantation through the largest
commonly used intra-uterine cannula, which has an internal diameter of 3.3mm. The
pacing system consists of a cylindrical micropacemaker containing the electronic
circuitry and a lithium power cell as well as a screw electrode plus its flexible lead.
The version illustrated in Figure 3-4 contains working electronics for pulse
generation powered by a primary cell. The design illustrated in Figure 3-5
incorporates an additional inductive coil and circuitry to allow recharging of a
secondary cell from outside the mother’s abdomen via a radio-frequency magnetic
field.
Figure 3-4 First Generation Fetal Pacemaker
A) Fetal micropacemaker within plastic insertion sheath (top); B) Pacemaker as deployed. Features
left to right: battery case of lithium cell, which functions as return electrode, glass sleeve for epoxy
encapsulation, printed circuit board with discrete surface-mount circuitry, flexible lead (0.075mm Pt-
20Ir with Parylene-C insulation), epoxy disk over welded joint, corkscrew electrode (0.150mm Ir with
Parylene-C insulation). (bottom)
65
Figure 3-5 CAD Model of the Fetal Pacemaker
Solid model of rechargeable micropacemaker incorporating circuitry.
The electronics for both the primary cell and rechargeable designs are protected
from body fluids by a thin-wall glass sleeve that is filled with a low permeability epoxy.
Polymeric encapsulation is not truly hermetic like the titanium cases commonly used
in pacemakers today, but it was a mainstay of cardiac pacemakers until the 1980s,
when battery life was typically limited to 18-24 months. Such lifetimes are possible
66
provided the diffusion path for water through the epoxy is long and the epoxy
completely surrounds and adheres to the circuitry without leaving voids in which
water can condense [42]. In the design illustrated in Figure 3-4 and Figure 3-5, the
glass sleeve provides a thin-wall form for the epoxy and a barrier to diffusion from
the lateral surfaces, so that the diffusion paths to the electronics are long and narrow.
The vacuum injection process for removing air bubbles and infiltrating the epoxy is
illustrated in Figure 3-6. After curing, the excess epoxy in the tube is removed by
cutting just in front of the glass sleeve and through a platinum rod that projects out
through the ferrite from the stimulus control capacitor (component C in Figure 3-7).
The subassembly consisting of the flexible lead (80% platinum – 20% iridium) and
cork-screw electrode (pure iridium) is fabricated separately and insulated with
vapor-deposited Parylene-C except for the exposed portion of the iridium electrode,
which is masked during deposition by embedding it in clay. Parylene has been used
to insulate chronically implanted electrodes because of it very high tensile strength,
flexibility and stability in vivo. The proximal end of the flexible lead is resistance
welded to the exposed cross-section of the platinum rod and reinforced with an
overcoat of the epoxy.
67
Figure 3-6 Injection Molding System Schematic
Epoxy injection molding system. The vacuum pulls air from the electronics and tubing through the slip
fit between the titanium battery case and the borosilicate glass sleeve. When the flow control clamp is
opened, epoxy flows slowly around the ferrite and copper coil and covers the printed circuit board
(PCB) and electronic components and battery terminals. Flow is stopped before the epoxy flows out
past the glass sleeve by releasing the vacuum. The epoxy is cured in place initially at room temperature
and then post-cured in an oven
The implantation and deployment scheme utilizes a thin-wall plastic sheath that
contains the pacemaker assembly during sterilization and surgical handling. The
epoxy reinforcing disk between the corkscrew electrode and the helical flexible lead
is wedged into the end of the plastic sheath, leaving only the electrode exposed. Under
ultrasound and fetoscopic guidance, the trocar and cannula are advanced from the
maternal abdomen, through the uterine wall and fetal chest wall, until it abuts the
68
fetal heart. The trocar is then removed and the plastic sheath assembly is inserted in
its place. The fetal surgeon affixes the electrode to the myocardium by pushing and
rotating the sheath so as to embed the protruding corkscrew electrode into the
myocardium. The pacemaker continuously generates stimulus pulses, which pass
into the myocardium from the electrode and return to the reference electrode
(exposed back end of the lithium cell casing) via a fenestration in the plastic sheath.
Adequate ventricular muscle capture can be assessed by ultrasonic imaging before
the pacemaker is released from the sheath. Release is accomplished by a push rod
(not illustrated) that advances the pacemaker assembly forward and therefore
releases the epoxy disk out of the end of the plastic sheath as the sheath is withdrawn
[43], leaving the pacemaker and unfurled helical lead lying in the chest of the fetus at
a slight distance from the epicardial electrode.
69
Figure 3-7 Pacemaker Circuitry Schematic
Schematic diagram of relaxation oscillator (center top) with typical output pulse (insert below), with
optional bypass diode Db for recharging through electrodes (dark grey box on left) and optional
recharging circuitry (light grey box at right) for transcutaneous inductive recharging from external coil
L1. See text for description of individual components. Insert below shows typical exponential cathodal
pulse from corkscrew Ir electrode (open arrow) with typical values for Ce after activation and Re in
myocardium.
Electronic Circuitry
The main challenge was to obtain the longest possible period of effective pacing
within the limits of a tiny, off-the-shelf lithium cell (Quallion QL0003I, Sylmar, CA)
70
and conventional surface-mount components. We achieved this by using a relaxation
oscillator, an old design for pacemakers [44] that requires a single active component
and delivers virtually all of its consumed power as stimulation pulses (schematic
diagram and output waveform illustrated in Figure 3-7). Output capacitor C is
charged through resistor RC until its voltage reaches the threshold of the
programmable unijunction transistor (UJT) as defined by voltage divider R1 and R2.
The UJT then switches to a low impedance state that allows all of the charge
accumulated on C to discharge through the electrode (an equivalent circuit is a
capacitor in series with a resistor). The UJT then switches back into a high impedance
state and begins another charge/discharge cycle whose rate is defined by time-
constant = RC*C. The component values in the schematic reflect the most likely design
based on the animal experiments described below.
Two approaches to charging the lithium cell are illustrated in grey boxes in the
schematic diagram in Figure 3-7. Adding a diode Db that bypasses the output
capacitor permits the rechargeable lithium cell to be maintained at full charge before
implantation by connecting the output electrodes to an external recharging circuit.
This recharging circuit can be powered by a relatively large primary cell that can be
incorporated into the sterile packaging. Alternatively, it would be far more desirable
to recharge the lithium cell repeatedly, including after implantation. This can be
accomplished by incorporating the inductive coil L2 and tuning capacitor Ct
illustrated in the schematic, which functions as a very weakly coupled transformer in
71
conjunction with an external recharging coil L1. Because the coupling coefficient can
vary widely depending on the orientation of the fetus, additional components need to
be incorporated to avoid damage to the lithium cell from excessive voltage (zener
diode Z1) or recharging current (limiting resistor Rr).
Figure 3-8 Pacing Rate Over Time
Stimulus rate (beats per minute) vs. hours of continuous pacing output for a high output design
(6.38µC @ 2.9Vcompliance, 130bpm at full charge). The initial rate increase is a conditioning effect in
the UJT when its threshold is set near Vs; it does not recur when recharged during continuous output.
The rate at which the pacemaker battery will discharge is reasonably predictable
(see below), but the effectiveness of the recharging process is variable due to the
frequent changes in the orientation of the fetus in the womb. The battery voltage
varies from approximately 3.0 to 4.0V depending on the state of charge, which causes
72
a small but readily detectable change in the output pulse rate (Figure 3-8). The brief
but relatively intense stimulus pulses in the fetus generate widespread stimulus
artifact signals that are easily recorded via skin electrodes on the mother’s abdomen.
We have modeled a proposed recharging scheme based on a 40cm diameter
transmitting coil operating at safe power levels in the 6.78MHz ISM band. Recharging
the Quallion cell will take less than 2 hours at its maximal recommended rate of
1.5mA/h, provided the coil in the implanted pacemaker is less than 18.7cm from the
plane of the transmitting coil on the mother’s abdomen and it is oriented
approximately parallel to its flux lines. Due to the ability to predict the state of charge
from the pacing rate, monitoring the maternal stimulus artifact rate throughout the
recharging process will verify effective recharging.
Electrode Properties
Stimulus efficacy depends primarily on charge density of the pulse in the tissue
immediately surrounding the active myocardial electrode. This is given by the
integral of the stimulus current waveform over time divided by the cross-sectional
area of the tissue. Stimulus efficacy depends secondarily on the rate at which that
charge is delivered (peak current). It is desirable to deliver the charge as rapidly as
possible, but this is limited by the available compliance voltage and the complex
impedance represented by the metal-electrolyte interfaces of the electrodes and the
intervening myocardial tissue: Istim = Vcompliance / (Zelectrode + Rtissue). With
73
smaller electrodes, the density of the charge they emit increases while the impedance
also increases. In the experiments described below, we tested two diameters of
corkscrew electrode (1.3 and 2.2mm) wound from 0.15mm diameter iridium wire
and sharpened at the tip using an abrasive wheel. The effect of the number of coils
that were electrically exposed was also tested. This was controlled by masking the
bare iridium wire during deposition of Parylene-C insulation [45].
At low voltages below the limits where unsafe redox reactions can occur, the
impedance can be modeled approximately as a resistance in series with a frequency-
dependent capacitance (Re and Ce). Ce depends on the electrochemical properties of
the metal electrode in the body fluids. We have chosen pure iridium for the electrode,
both for its high mechanical strength (Young’s modulus = 528GPa) and for its ability
to be “activated” by growing a conductive, porous oxide on its surface. [46] This
activation was accomplished by cyclic voltammetry, in which a potentiostatically
controlled voltage ramp (0.5V/s) was applied to the electrode in phosphate-buffered
physiological saline while monitoring the resulting current flow. As the oxide layer
develops, its capacity to store and release charge increases greatly, as shown by the
increasing area enclosed in successive cycles of the voltammogram in Figure 3-9, A.
This is reflected in impedance spectroscopy results in normal saline (Figure 3-9, B),
which shows substantial reductions in the impedance over time, particularly at lower
frequencies. On the basis of these results, we adopted 20 minutes as the standard
activation time for the electrodes tested in vivo.
74
Figure 3-9 Cyclic Voltammetry
Top: typical cyclic voltammetry curves (current on abscissa vs. applied voltage with respect to
saturated calomel electrode in phosphate buffered physiological saline) for ±0.5V/s ramp between
cathodal (1) and anodal (6) electrolysis points showing progressively larger enclosed charge; Bottom:
impedance (solid lines, left ordinate) and phase (dotted lines, right ordinate) for various sinusoidal
frequencies as a function of activation time (abscissa, minutes) for 1.3mm diameter Ir corkscrew
electrode with 4 coils (2mm) exposed length in saline.
75
Results
Animal work was approved by and performed in accordance with guidelines of
the University of Southern California's Animal Care and Use Committee (IACUC,
protocol 11686), which adheres to the Guide in the Care and Use of Laboratory
Animals established by the U.S. National Academy of Sciences.
In Vivo Percutaneous Implantation
The tools and techniques for percutaneous device implantation were tested using
an in vivo rabbit model. A non-functional device with identical dimensions and
mechanical features of our fetal micropacemaker was implanted into an adult rabbit
under general anesthesia, using a conventional cannula (3.8mm o.d. x 3.3mm i.d.) and
the insertion sheath illustrated in Figure 3-4. The device was implanted into the
actively beating rabbit myocardium from a subxyphoid approach, using
transcutaneous ultrasound guidance and without the need for surgical incisions. A
small amount of axial force on the insertion tool stabilized the motion of the ventrical
at the tip of the corkscrew electrode as axial rotation implanted it into the
myocardium in a few seconds. The animal was subsequently euthanized and
explored surgically. We were able to confirm appropriate placement of the electrode
in the myocardium with a favorable and safe position of the pacemaker and lead in
the rabbit thorax (Figure 3-10). This demonstrates the feasibility of the strategy and
76
instrumentation. Further mechanical implantation tests will be performed on a fetal
sheep model, a more appropriate model of our clinical goals.
Figure 3-10 Percutaneous implantation site at post mortem exploration
Explantation shows the mock pacemaker is anchored in the myocardium.
In Vivo Pacing Threshold
A set of activated iridium electrodes with varying coil diameters and exposure
lengths were prepared and affixed to the end of a handle similar to the proposed
insertion sheath. They were tested in three anesthetized adult rabbits weighing
approximately 3kgs whose beating hearts were exposed via thoracotomy and then
sacrificed with a barbiturate overdose. The electrodes were screwed into the
77
myocardium at various locations (Figure 3-11, A) while monitoring the surface ECG.
Pacing threshold was defined as the minimal value that produced ventricular
myocardial capture (demonstrated by the presence of premature ventricular
contractions on the ECG). Electrode impedance was measured in situ using a 1 kHz,
10µA sinewave. Stimulation pulses were generated by a conventional clinical
analyzer for pacemaker leads (Merlin Interrogator from St. Jude Medical, Sylmar, CA)
that allowed independent control of square wave voltage and duration to determine
threshold for capture. The stimulus charge was estimated from the measured
electrode impedance (charge = duration x voltage/impedance). We also used a
custom-built instrument that contained the UJT and associated circuitry for our fetal
micropacemaker but with the ability to systematically vary C, Rc and Vs to determine
thresholds for its exponential output pulse.
78
Figure 3-11 Strength-Duration Curves for Various Implantation Sites
Top: insertion locations for data in Table 1; Middle: threshold strength-duration curves for square-
wave stimuli; Bottom: Comparison of threshold charge (red lines; right ordinate) vs. pulse duration for
square waves (black dots) and exponential pulses from relaxation oscillator (vertical red bar at time
constant = 0.375ms) for pacing at Spot #1.
79
The strength-duration curves for three insertions of the 1.2mm diameter
corkscrew electrode using a conventional square-pulse generator (Merlin
Interrogator) are plotted in Figure 3-11, B. The relaxation oscillator produces an
exponentially declining stimulus (insert in Figure 3-7) that does not provide explicit
control of current or duration, which will change depending on the actual impedance
of the electrode. It does, however, provide precise control of the pacing charge (and
hence power consumption) by varying capacitor C. Importantly, the charge required
to pace was comparable to that of the square wave; see bracketed red line in Figure
3-11, C for one placement of a 1.2mm coil electrode. The chronaxie for a conventional
square-wave stimulus is approximately 0.5ms and its threshold charge lies between
the total charge of the exponential pulse at infinity and the charge delivered at one
time constant (τ=0.375ms). Note that only relatively coarse steps (factors of 2) were
used for C, corresponding to the limited range of medical grade capacitors available
in the 0402 surface mount package required for the implant.
Table 3-1 provides a complete set of data for the set of electrodes and placements
illustrated in Figure 3-11, A; similar but less complete data were obtained from the
other two animals (4 placements in each animal, threshold charge 0.53 – 2.4µC). The
electrode impedances and thresholds were somewhat variable but all within a useful
range except for placement in Spot #3. The electrode did not penetrate the
myocardium well at this location and the right ventricular wall at this location
appeared to twist as the electrode was inserted due to the thinner myocardium. After
80
electrode removal, the location appeared erythematous, unlike other insertions. The
main factors affecting capture thresholds appeared to be the sharpness of the beveled
tip of a given electrode and the number of different times it was inserted. In general
the smaller diameter coils (1.2mm used exclusively in this animal) were more stable
mechanically and easier to physically insert by applying gentle pressure and rotating
the handle axially, mimicking the anticipated insertion technique when operating
through the uterine cannula. Importantly, due to the flexibility of the electrode, there
was no difficulty inserting the smaller corkscrew electrode even when oriented at a
45-60º oblique angle to the myocardial surface, as may well occur when the device is
implanted in a human fetus in utero. Impedances and thresholds with the larger coils
(2.3mm diameter) with similar lengths of exposed electrode were similar to those of
the small coils but they were physically more difficult to implant, particularly for
oblique insertions.
81
Electrode #1 #1 #2 #2 #3 #3
Electrode
coil diameter
1.3 mm 1.3 mm 1.3 mm 1.3 mm 1.3 mm 1.3 mm
Exposed 3 turns 3 turns 3 turns 3 turns 4.5 turns 4.5 turns
Spot Spot #1 Spot #2 Spot #3 Spot #4 Spot #5 Spot #6
Location LV apex
Lateral
LV apex
High RV Diaphragmatic
RV free
wall
LV apex
(infra)
Average
impedance Ω
408 415 362 428 313 312
Square-wave
threshold
(1.3 ms dur.)
0.5V 0.75V 2.75V 1.25V 0.5V 0.5V
Square-wave
threshold
(0.5 ms dur.)
0.75V 1V 3.75V 1.5V 0.75V 1V
Exponential
threshold
(µC total)
1.03 1.37 >2.4 2.40 1.03 0.79
Table 3-1 Summary of Electrode Performance Based on Location
Electrode and pacing data from placements illustrated in Figure 3-11, A.
Power Optimization
Maximizing battery life depends on minimizing the power consumed by the
stimulus pulses themselves and by the biasing resistors for the UJT (voltage divider
R1-R2 in the schematic in Figure 3-7). Because the output of the pacemaker is
determined by the fixed component values, the stimulus charge must include a
reasonable safety margin over the measured capture thresholds to assure reliable
pacing. The charge is the product of the capacitor value and the voltage range
between peak charging (the threshold voltage for discharge through the UJT and
82
complete discharge (approximately zero volts). When the UJT threshold is set near
the maximal voltage to which the capacitor can be charged (i.e. the supply voltage Vs
from the lithium cell), the UJT will not trigger reliably unless the values of R1 and R2
are set fairly low (<200K), resulting in substantial continuous current flow through
the voltage divider. When Vthreshold = 0.8Vs, R1+R2 = 300K, resulting in 10µA,
comparable to the continuous current needed to recharge output capacitor C. This
problem can be minimized by reducing the threshold voltage to 0.5Vs, in which case
R1 = R2 = 2MΩ and biasing current is <1µA. In order to maintain stimulus charge with
a smaller voltage excursion, C must be increased accordingly. This has the effect of
increasing the time constant of the stimulation pulse (product of C and tissue
resistance Re), but the in vivo pacing data (above) indicate that these pulses are still
shorter than the chronaxie for the heart muscle, so should produce effective and
efficient pacing.
For the component values indicated in the schematic in Figure 4, the output pulse
rate is 130 beats per minute with a charge of 0.73 µC over a time constant of 0.375ms.
The charging current is 1.58µA and the biasing current is 0.9µA; leakage current
through the UJT is negligible. The 3mAh battery should last 50d on a single charge.
These values, however, represent the best case conditions observed in the in vivo
experiments (e.g. Figure 3-11, C). A safety margin of at least a factor of 2:1 for the
stimulus charge compared to the capture threshold would be prudent (see Table 3-1)
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and more may be indicated as a result of experiments in fetal sheep that remain to be
performed.
Discussion
While the results to date have been encouraging, a number of steps remain before
this device is ready for use in human fetuses.
Implanting and Pacing a Fetal Heart
The adult rabbits used in the experiments reported here have hearts similar in
size to a 28-week human fetus but the chest wall anatomy and the inflated chest are
significant differences of this model from that of human fetal implantation. The next
set of animal implantations will be in a fetal sheep model, which not only provides a
more physiologically appropriate model, but also allows chronic instrumentation
(including femoral venous and arterial lines and fetal sheep skin electrodes for fetal
ECG monitoring) without inducing premature labor [47, 48]. A functional
micropacemaker can be implanted percutaneously and left in place for the remainder
of the pregnancy so that its efficacy, orientation and tissue encapsulation can be
assessed. The pulse parameters required to pace the normal adult rabbit heart
(Figure 3-11 & Table 3-1) or a normal fetal sheep heart may differ from those
required in a hydropic human fetus, but data from a study by Assad et al. suggests
84
that the differences may not be large. In that study, a strength duration curve was
created using a different electrode design (T-bar shape without screw) in a 25-week
hydropic human fetus at the time of implant and on the first postoperative day (POD)
[36]. The patient died by 36 hours after implantation of the lead (which was
connected to a pacemaker in the maternal abdominal wall) and the electrode design
was different from our screw mechanism, but data from their strength-duration curve
suggest a pacing threshold and chronaxie similar to our adult rabbit studies.
While our current electronic configuration results in a pacing rate of 130bpm, we
anticipate that a rate of 100bpm would be used for human fetal application. This value
is based on clinical experience with newborn pacing, and we therefore do not
anticipate any negative cardiovascular effects from this increase in rate from the fetal
heart block baseline rate (expected to be 40-60bpm). After the electrode screw is
implanted in either a fetal sheep model or a human fetus, proper placement in the
myocardium will be confirmed by determination of adequate pacing prior to release
of the pacemaker from the insertion sheath. This will allow for the possibility of
multiple screw insertion attempts prior to ultimate device placement, and will be
particularly important when the fetal orientation results in more technically
challenging implantations.
85
Extending Functional Life
The currently available rechargeable lithium cell is expected to support pacing for
15-20 days on a single charge. For fetuses that develop hydrops fetalis relatively late
in gestation, continuous pacing for 2-3 weeks may be sufficient to resolve the hydrops
and permit cesarean delivery with a good chance of survival. For other cases, however,
we expect pacing will be required for a longer period. Primary lithium cell chemistry
generally provides about three times the charge density of rechargeable chemistry,
albeit at a somewhat lower voltage. A fetal pacemaker with the same dimensions
could be constructed from a custom primary cell. If it were still near full charge at the
time of implant, the 6-9 week device longevity would likely be sufficient for most, if
not all, fetuses with heart block. Alternatively, the rechargeable lithium cell could be
repeatedly recharged using the inductive coupling system described above. The
external RF transmitter / coil could be used in a clinic or in-patient environment
where it would be powered from an 110VAC receptacle. The recharging range would
depend on the coil diameter and field strength that can be achieved without excessive
heating of the coil or nearby skin. The magnetic flux density required to reach a
sufficient voltage in the implant coil to recharge the battery depends on the distance
between the two weakly coupled coils and their relative orientation [49]. These
factors can be observed using ultrasound but not controlled given the mobility of the
fetus in the uterus. We anticipate that recharging will be performed under medical
supervision and with monitoring of the actual battery charge as predicted by the
86
pacing rate measured from the electrical artifact recordable on the mother’s abdomen
(see Figure 3-5).
It may be desirable to use percutaneously implanted cardiac pacemakers in young
children and even adults for which surgical implantation and intravenous passage of
the electrical lead is undesirable. In most such applications, a functional life of years
would be required. Substantial increases in the size of the pacemaker and its battery
would likely be possible, which could substantially increase the recharging interval.
It would then be feasible to use conventional hermetic packaging and to add
bidirectional telemetry and digital control functions such as stimulus
programmability. The fixed rate of the simple relaxation oscillator employed in the
prototype presented here is not ideal but should be adequate for the last trimester of
pregnancy. Perhaps the biggest challenge would be a flexible lead that could
accommodate the small but continuous movement between the electrode in the
myocardium and the pacemaker package in the chest.
Device Longevity and Biocompatibility
The fetal pacemaker will likely become ineffective shortly after delivery. The rapid
expansion of the lungs and chest will certainly put stress on the flexible lead and the
anchoring of the electrode in the myocardium. The open spiral of the flexible lead is
not designed to accommodate large stresses or continuous flexing from respiratory
87
and cardiac motion and the lithium cell will soon run down. Thus, we assume that a
small, conventional demand pacemaker will be surgically implanted soon after birth.
The question arises as to whether and when a non-functional fetal pacemaker
needs to be removed from the infant. The materials that form the outside surfaces of
the fetal pacemaker include titanium battery case, iridium electrode, Parylene-
insulated lead, borosilicate glass sleeve, and medical-grade epoxy. All have a long
history of use in chronically implanted medical devices, so we anticipate successful
biocompatibility testing, which would be required on the completed device.
Accelerated life-testing of the non-hermetic packaging at elevated temperature and
pressure in saline will be required to demonstrate that the electronic circuit
continues to function for the required period of at least 3 months. These tests can be
extended to demonstrate whether there are degradation modes of the inactive device
that could pose a danger, such as swelling and disruption of the epoxy package.
Clinical Translation
It is estimated that approximately 500 pregnancies in the United States are
affected by fetal heart block annually and may be candidates for this device.[38] In
view of this relatively small market, we have applied for and been awarded a
Humanitarian Use Device [HUD; 21 CFR 814(h)] designation by the US Food & Drug
Administration (FDA). Under this designation, first-in-human studies after the
88
completion of the ongoing development and testing can be pursued according to a
Humanitarian Device Exemption (HDE), rather than standard Investigational Device
Exemption (IDE) processes. The HDE pathway allows a product to be marketed with
a demonstration of safety rather than efficacy and with reduced Good Manufacturing
Practices (GMP) requirements compared to a Premarket Approved (PMA) product.
If our percutaneously implantable micropacemaker proves successful, it will
provide an extremely effective treatment option for a population of fetuses that
would either die in utero or require premature delivery with all of its comorbid
consequences. The fetal market is too small to attract major investment by industry,
so we have identified a way to address the key anatomical, physiological and surgical
issues while using off-the-shelf electronic components. Commercial organizations
(including Medtronic, Inc., Minneapolis, MN), have recognized the need for smaller
pacing devices and are developing “leadless” pacemakers for transvenous
implantation. These models, however, are designed for endovascular use and do not
have the active fixation mechanisms necessary for fetal use, nor the flexible lead
systems required for transthoracic implantation. Once available, our class of
injectable epicardial devices could be extended to newborns, infants, and even adults
with limited venous access and/or contraindications to open surgery.[50] As battery
technology inevitably improves, this technology could replace standard single-
chamber pacemaker techniques with implantation of the entire pacing system into
the patient’s thorax via a minimally invasive technique. In this time of expanding
89
micro- and nano-technology, we see our percutaneously implantable
micropacemaker as the natural evolution of pacemakers and anticipate a much wider
application of this and related microdevices.
Acknowledgments
This research was funded by the Southern California Clinical and Translational
Science Institute and by the Robert E. and May R. Wright Foundation. The authors
thank consultant Glen Griffith for feasibility analysis of the inductive recharging
scheme, and thank Dr. Erlinda Kirkman for providing veterinarian assistance for the
animal models used in this report.
90
Chapter 4: Electronics, Pacing Thresholds, and Power Budget
Adriana Nicholson, Ramen Chmait, Yaniv Bar-Cohen, Kaihui Zheng, and Gerald E. Loeb
© 2012 IEEE. Reprinted from:
Nicholson, A., Chmait, R., Bar-Cohen, Y., Zheng, K. (2012). Percutaneously
Injectable Fetal Pacemaker: Electronics, Pacing Thresholds, and Power Budget.
Engineering in Medicine and Biology Society (EMBC), 2012 Annual Conference
of the IEEE EMBS San Diego, CA
Preface
This publication details a power budget analysis for the pacemaker. It first shows
the requirements of output of the device in order to elicit effective electrical
stimulation of the myocardium in an animal model. Based on the information learned
about stimulation thresholds, current drain is optimized for a minimum charge
delivery and minimum bias current. The analysis shows that recharging the device
will be necessary in order to meet the design requirement lifetime of three months.
Current design of the device will require multiple recharging cycles, but it is still
desirable to minimize the amount of drain current in order to maximize the amount
of time the pacemaker can operate between recharging cycles. Minimal current drain
is difficult to attain because of peculiarities of the programmable unijunction
transistor (PUT) and the way it behaves in the relaxation oscillator. Those
considerations are included here as well.
91
My contribution to this research includes the analysis of strength duration
thresholds from the preliminary animal studies and using that information to
optimize the amount of power draw from the battery. A better understanding of
stimulation thresholds allows the design of the pacemaker to incorporate an
adequate safety factor of charge injection without using excessive charge.
After analyzing the threshold information from animal experiments, it became
clear that in a best case electrode placement (that is a placement with a low threshold
for capture), the amount of energy needed to capture the heart muscle was about 0.73
µC with a time constant of 0.375ms. With this information in mind, the components
of the relaxation oscillator could be chosen so that the current needed to charge the
output capacitor was 1.58 µA and the current needed to bias the oscillator’s transistor
was below 1 µA. Given these drain currents, a 3 mAh capacity Li cell such as the one
employed in the design of the pacemaker could reliably pace for up to 50 days without
recharging.
Incorporating a safety margin in the threshold is a high priority of the design, as
it would be unfavorable to implant a device only to find that the threshold of the
electrode tissue interface is too high to capture at a given location. Unlike modern
pacemakers, the stimulation energy cannot be adjusted after fabrication. The
implanting physician in this case would need to either implant a second device or
remove and relocate the device already implanted. Therefore, the drain current in the
92
final design will be higher than previously stated and the battery will be depleted in
a shorter amount of time.
Abstract
We are developing a cardiac pacemaker that is designed to be implanted
percutaneously into a fetus to treat complete heart block and consequent hydrops
fetalis, which is otherwise fatal. One of the most significant considerations for this
device is the technical challenges presented by the battery and charging system. The
size of the device is limited to about 3 mm in diameter; batteries on this scale have
very small charge capacities. The smaller capacity means that the device needs to be
designed so that it uses as little current as possible and so that its battery can be
recharged wirelessly. We determined the pacing thresholds for a simple relaxation
oscillator that can be assembled from discrete, surface mount components and
analyzed the power consumption of the device given different electrode
configurations and stimulus parameters. An inductive recharging system will be
required for some patients; it is feasible within the package constraints and under
development.
93
Introduction
Complete heart block in the fetus is a life-threatening emergency with no effective
treatment options beyond watchful waiting [28, 29]. Fetal bradycardia due to heart
block can progress in utero, and for more than a quarter of these fetuses may result
in hydrops fetalis [30, 31]. Once hydrops fetalis develops, if the fetus cannot be
delivered due to prematurity or other clinical concerns, fetal demise is nearly
inevitable [29]. Pharmacological and immunological therapies delivered systemically
via the mother have not been effective in clinical trials [32-34].
When a newborn, child or adult presents with symptomatic complete heart block,
treatment usually consists of implantation of a pacemaker to ensure an adequate
heart rate. With appropriate pacing, these patients are usually asymptomatic with an
excellent prognosis. Similar benefits would be expected from pacing a fetus with
complete heart block, theoretically allowing resolution of hydrops in 1-2 weeks and
permitting an otherwise normal gestation. A conventional pacemaker would then be
implanted at delivery. Over the last two decades, several investigators have
attempted to place pacemakers in a fetus. [35-37] To date, however, there have been
no survivors of fetal pacing. Previous approaches have relied on the placement of a
pacing wire on the fetal heart with an extra-uterine pulse generator implanted in the
mother. This has inevitably failed because of lead dislodgement due to fetal
movement.
94
We have designed a single-chamber pacing system that is self-contained and can
be completely implanted in the fetus without exteriorized leads, thereby permitting
subsequent fetal movement without risk of dislodgement of the electrodes (Figure
4-1). Such a design is now possible because of significant developments in medical
device miniaturization and advances in fetal surgical intervention, allowing the
pacing system to be percutaneously deployed through the maternal abdomen under
ultrasound guidance.
Figure 4-1 CAD Model
Solidworks model of the rechargeable micropacemaker within plastic insertion sheath. The corkscrew
electrode at the left is made from pure iridium that has been activated to achieve a low interface
impedance with tissue by growing a layer of iridium oxide on its surface. Other components from left
to right include epoxy disk wedged into end of transparent insertion sheath, flexible helical lead
compressed into sheath, inductive recharging coil wound over hollow ferrite core, printed circuit
board with surface mounted components, and rechargeable lithium cell with hermetic case functioning
as reference electrode. [51]
Diameter
of 3.3 mm
95
Device Design
The fetal pacemaker is designed to be a simple, compact, low-power device that
generates brief electrical stimuli that initiate ventricular contractions at a
physiological rate. Relaxation oscillators have high reliability and draw little power
while they are operating, making them well suited for this application. The relaxation
oscillator for this device uses a programmable unijunction transistor (PUT, Figure
4-2), paired with a capacitor and a battery. The battery charges the capacitor while
the PUT remains off. Once the capacitor exceeds a gate voltage on the PUT, the PUT
turns on and creates a current path that discharges the capacitor through the heart
tissue. Once the current in that path drops to a certain level, the PUT reverts back to
a high resistance state and the capacitor can begin charging again. The gate voltage
for the PUT is set by biasing resistors R1 & R2. The stimulus charge depends on the
product of capacitor C and the gate voltage. The stimulus rate depends on the time
constant of capacitor C and charging resistor RC. All of these component values must
be optimized to achieve the desired combination of reliable pacing and long battery
life. Once it is finalized, the system can be built on a printed circuit with surface-
mounted components, making the profile as slim as possible.
The battery (actually a single, rechargeable lithium cell) to be used was chosen for
its size, about 3 mm in diameter, and its hermetic packaging. It is manufactured by
Quallion and has a nominal capacity of 3mAh. The mean discharge rate and operating
temperature for this application are well within normal operating range for this cell,
96
so it is expected that it will perform according to the nominal specifications from the
manufacturer.
Methods
To understand better the pacing threshold of a fetal heart, we paced three
anesthetized adult rabbits whose beating hearts were exposed via thoracotomy and
then sacrificed with a barbiturate overdose. The electrodes were screwed into the
myocardium at various locations while monitoring the surface ECG. Pacing threshold
was defined as the minimal value that produced ventricular myocardial capture
(demonstrated by the presence of premature ventricular contractions on the ECG).
Stimulation pulses were generated by a conventional clinical analyzer for pacemaker
leads (Merlin Interrogator from St. Jude Medical, Sylmar, CA) that allowed
independent control of square wave voltage and duration to determine threshold for
capture. The stimulus charge was estimated from the measured electrode impedance
(charge = duration x voltage/impedance). We also used a custom-built instrument
that contained the UJT and associated circuitry for our fetal micropacemaker but with
the ability to systematically vary C, Rc and Vs to determine thresholds for its
exponential output pulse.
After animal studies were completed, a power budget was developed with the goal
of minimizing the amount of energy being used by the system for any purpose other
97
than directly stimulating the heart. The budget was compared to previous benchtop
tests of battery lifetime to verify the device was acting predictably.
The analysis of battery life was performed by summing the average amount of
current (Coulombs per second) through each pathway from the battery to the tissue
or ground. There are three paths for current to flow. First, there are the passive
elements of the relaxation oscillator, the gate voltage divider of R1 and R2. The
current through this part of the pacemaker is constant and does not contribute to the
output stimuli. The second path is through RC and C during the charging phase, the
current of which was found by taking the time average of the amount of charge over
each cycle. Lastly, the current through RC during the discharge phase when the PUT
grounds this resistor was found and averaged over the time when the PUT is closed.
Battery voltage was set to 3.6 V (it will actually vary between 4.0 and 3.0
depending on charging status) and pulse rate set to 130 beats per minute. All other
parameters were varied.
98
Test Results
Pacing Threshold Measurement
The strength-duration curve for an insertion of a 1.2mm diameter corkscrew
electrode using a conventional square-pulse generator (Merlin Interrogator) is
plotted in Figure 4-3. The relaxation oscillator produces an exponentially declining
stimulus (insert in Figure 4-2) that does not provide explicit control of current or
duration but does provide precise control of the pacing charge (and hence power
consumption) by varying capacitor C. Importantly, the charge required to pace was
comparable to that of the square wave; see bracketed red line in Figure 4-3 for one
placement of a 1.2mm coil electrode.
99
Figure 4-2 Pacemaker Schematic
Schematic diagram of relaxation oscillator (center top) with typical output pulse (insert below), with
optional bypass diode Db for recharging through electrodes (dark grey box on left) and optional
recharging circuitry (light grey box at right) for transcutaneous inductive recharging from external coil
L1, including tuning capacitor Ct, voltage limiting zener diode Z1 and charge current limiting resistor
Rr. Insert below shows typical exponential cathodal pulse from corkscrew Ir electrode (open arrow)
with typical values for Ce after activation and Re in myocardium.
The chronaxie for a conventional square-wave stimulus is approximately 0.5ms
and its threshold charge lies between the total charge of the exponential pulse at
infinity and the charge delivered at one time constant (τ=0.375ms). Note that only
relatively coarse steps (factors of 2) were used for C, corresponding to the limited
range of medical grade capacitors available in the 0402 surface mount package
required for the implant.
100
Figure 4-3 Strength-Duration and Charge
Schematic Comparison of threshold charge (red lines; right ordinate) vs. pulse duration for square
waves (black dots) and exponential pulses from relaxation oscillator (vertical red bar at time constant
= 0.375ms) for pacing at the Left Ventricular Apex
Figure 4-4 Pacing Rate Over Battery Discharge Time
Stimulus rate (beats per minute) vs. hours of continuous pacing output for a high output design
(6.38µC @ 2.9Vcompliance, 130bpm at full charge). The initial rate increase is a conditioning effect in
the UJT when its threshold is set near Vs; it does not recur when recharged during continuous output.
101
Power Budget
Trial 1 was modeled after a test performed on battery life on the benchtop. The
mean current delivered by the pulsing system is about 10 µC/second. The charge of
one pulse is about 4.7 µC. The pulse rate is about 130 beats per minute; the pulse
duration, from the peak of the pulse to 33% of the peak, is 0.6 msec. These parameters
with a 3mAh capacity battery predict that the battery will last 9 days (Table 4-1). A
physical trial of the battery life was done with identical parameters. The battery
voltage varies from approximately 3.0 to 4.0V depending on the state of charge, which
causes a small but readily detectable change in the output pulse rate when the device
is nearing full discharge (Figure 4-4). The pacemaker was fully charged and then
allowed to completely discharge while rate was observed, and showed a lifetime of
about 10 days.
Charge delivered over time (C/s = A)
Trial 1:
Benchtop test
configuration
Trial 2: Higher
resistance through
passive elements
Trial 3: Variability
among PUTs
R1 and R2 2.81 uA 0.9 uA 5.19 uA
Exponential
decay pulse
10.1 uA 1.58 uA 2.22 uA
RC 26 nA 1.93 nA 2.93 nA
Total
Discharge
Current
12.9 uA 2.48 uA 7.41 uA
Days 9.7 days 50.3 days 16.9 days
Table 4-1 Power Budget Over Different Conditions
Variability in operation among the transistors (PUTs) leads to the average operation capabilities being
lowered.
102
Animal tests demonstrated that much smaller charge pulses than those used
during initial development could usually stimulate the heart. As we reduce the
amount of current drawn by the active pacing part of the device, the current drawn
by the passive part of the device starts to have more influence on the battery life. The
actual values of R1 & R2 are limited by instability in the PUT that occurs when the
threshold is near the available power supply voltage, so reducing the stimulus charge
by reducing the voltage excursion on C (which is equal to the threshold voltage on the
PUT) provides an opportunity to reduce passive power consumption
disproportionately. The animal tests suggested that a pulse of 0.73 µC would provide
an adequate safety margin to assure reliable pacing for most electrode placements in
the ventricular myocardium. Trial 2 of the power budget shows that raising values
for R1 and R2 has a dramatic effect on the predicted battery lifetime, in addition to
the smaller charge required for the actual stimulus pulses. With these changes, the
predicted lifetime of the device is 50 days, which would mean some patients will not
require recharging.
One limitation to raising the values for R1 and R2 is that there was considerable
variability among samples of the programmable unijunction transistors. Therefore,
the next budget (Trial 3) was calculated to achieve an average lifetime given a
variability in PUTs. In order to keep the same gate voltage for the three tested PUTS,
lower values for R1 were used, yielding a higher gate voltage (1.80V instead of 1.28V)
for this trial. The average PUT would give a substantial reduction in battery life. When
103
we triple the discharge current in Trial 3, we get a third of the battery life as compared
to Trial 2. This shows that the values for the resistors R1 and R2 still have a significant
effect on the power consumption, and steps should be taken to find PUTs that are less
variable with more optimal characteristics.
Discussion
The data from the animal experiments and bench-testing of the component values
will be combined into a final design for fabrication and testing. The capacitor and the
gate voltage dictate how much charge is delivered to the tissue for each pulse. We are
aiming for a charge of about two times the average threshold. Such a safety factor is
necessary because the stimulus parameters cannot be adjusted after implantation.
Fortunately, pacing capture can be confirmed while the pacemaker is still attached to
its insertion sheath, permitting the electrode to be relocated in the myocardium if
necessary before release of the device into the chest of the fetus.
We estimate that the final design will require a mean battery current of about 4
µA, so the 3mAh lithium cell should last about 34 days on a single charge. The cell can
be kept charged until the time of implantation by applying recharging current
through the output electrodes via the proposed bypass diode Db. Fetuses that develop
hydrops early or cannot be delivered prematurely will require more than 3 to 4 weeks
of pacing, however. An inductive recharging circuit will need to be integrated into the
104
implant, as shown schematically in Figure 4-1. Future work will involve designing a
primary coil and class E RF oscillator for transcutaneous charging.
Acknowledgments
This research was funded by the Southern California Clinical and Translational
Science Institute of the and by the Robert E. and May R. Wright Foundation. The
authors thank consultant Glen Griffith for feasibility analysis of the inductive
recharging scheme.
105
Chapter 5: Development and Testing of a Closed Loop Recharging
System
Adriana N. Vest, Kaihui Zheng
Preface
The pacemaker will need to operate longer than the drain current and battery
capacity will support in the majority of treatment cases. Therefore, the battery will
need to be recharged repeatedly over the functional life of the implanted device, using
transcutaneous RF inductive coupling. The strength of that coupling is highly variable
because of the movement of the fetus, so the recharging process must be carefully
monitored and controlled to be sure that it is proceeding to completion without
overcharging or damaging the battery. A closed loop recharging system was
developed to address this need. The following work proves the feasibility of the
system on the benchtop, including design aspects to optimize performance.
Background
To determine when to charge and to monitor the charging process itself, we take
advantage of the change in pacing rate with battery voltage. Change in pacing rate
with primary supply voltage was used to decide when to replace the earliest cardiac
106
pacemakers circa 1960, which were based on a relaxation oscillator circuit similar to
what we have used in this device. As the battery's charge is drained, the battery
voltage decreases and results in a small but detectable change in the rate of output
pulses as the internal voltage drops in the transistor become a larger fraction of the
battery supply voltage to the circuit. This relationship can be seen in Figure 5-1. We
use the relationship between output pulse rate and supply voltage to monitor charge
level and recharging rate of the battery.
In order to control the charge level and recharging rate of the battery, the pacing
rate needs to be correlated with the true supply voltage. This is calibrated during
fabrication of the device. The supply voltage is varied while output pulse rates are
measured, creating a calibration curve (Figure 5-1). The stimulus artifact from the
fetal pacemaker is easily recorded from the abdomen of the mother, so the supply
voltage can be calculated from the pacing rate according to the stored calibration data.
From this voltage, it can be determined whether the lithium cell is within the
acceptable operational window, from 3.0 V to 4.0 V, and recharging can be initiated
or terminated. In addition to monitoring these upper and lower limits of charge, it is
important to limit the charging current, which varies depending on the strength of
coupling between the primary and secondary RF coils. The charging current can be
determined by knowing the internal resistance of the battery. From Ohm’s law, any
charging current that is applied to the battery will produce an added voltage
according to E = IR. For example, a 3V lithium cell with an internal resistance of 100
107
ohms that is being recharged with 1mA of current will actually generate 3.1V. By
accurately measuring the output pulse rate when the recharging field is turned on
and again when it is turned off, this voltage change and the actual charging current
can be computed. If that recharge current is excessive for the ratings of the battery,
then the field strength produced by the external transmitter is automatically reduced
until the rate change is consistent with the desired recharging current.
Figure 5-1 Relationship Between Output Pulse Rate and Supply Voltage
The voltage supplied by the lithium cell can be inferred from the pacing rate at any given time. During
recharging, the rate of charging could also be determined by comparing the pacing rate with the
charging current on, to that with the charging current off. The difference between the supply voltages
inferred from those rate reflects the IR drop induced by the recharging current passing through the
internal resistance of the lithium cell (and any additional resistance that may be added to the circuit
as described below).
108
Investigation Into Component Values
Component values tend to have a dramatic effect on the relationship between
pacing rate and supply voltage. Therefore, it is important to carefully investigate
these relationships. The turn-on voltage is determined by the voltage at the gate of
the PUT, which is set by the biasing resistors. The frequency of stimulation, also
known as the pacing rate and measured in beats per minute (BPM), is determined
primarily by the value of the output capacitor and the series resistor (RC) via which
it is charged to the turn-on voltage. Additional factors that affect the selection of
components are the characteristics of the PUT. The PUT in this application is being
used at the limits of its capabilities, at very low frequencies and voltages. Therefore it
will only operate with certain component ranges. The various relationships of these
components were considered by simulations and empirical breadboard evaluations
by testing different combinations.
Ratio Between Bias Resistors
It is helpful to understand first how the PUT works in the circuit to create an
astable condition and therefore oscillations. It is a combination of a PNP and NPN
transistor that is set up like a thyristor, except with the gate contact on the N-doped
segment as shown in Figure 5-2. When the top PN junction is forward biased, with a
higher voltage on the anode than the gate, current will flow and causes a negative
resistance effect. This is seen in Figure 5-3 on the downward slope that signifies a
109
decreasing voltage with increasing current. The anode of the device is pulled down to
about a diode drop from ground. In order to create oscillations, the charging resistor
RC must be small enough to supply enough current to raise the anode to VP, the peak
point, while charging the capacitor. Once VP is reached, anode voltage decreases as
current increases, which moves the operating point to the valley. It is the job of the
capacitor to supply the valley current IV. Once the capacitor is discharged, the
operating point resets back to the peak point.
The resistor RC must be small enough so that it can supply enough current to
overcome the peak at IP and get into an oscillation, but large enough so that it will
never supply the high valley current IV. In this application, operation occurs at the low
end of available frequency (2 Hz) and therefore RC’s upper limit is not of major
concern. Instead, the focus has been on the valley current. If the charging resistor is
too small and supplies IV, the operating point would never reset back to the high
resistance condition to the left of the peak point after the capacitor was discharged.
The valley current is a dynamic value that is a function of supply voltage and RG,
which is the gate impedance and is affected by the values of R1 and R2.
𝑅𝐺 =
𝑅 1 × 𝑅 2
𝑅 1 + 𝑅 2
Therefore, changes to RC, to the bias resistors R1 and R2, or to the supply voltage
will affect whether the device will oscillate.
110
Figure 5-2 Programmable Unijunction Transistor
(Left) Layout of components in a programmable unijunction transistor [52]. (Right) Valley current for
the device is a function of RG and supply voltage as seen in the datasheet [53].
Figure 5-3 Voltage Vs. Current in PUT
Relationship between voltage and current in the PUT [53].
In addition to understanding when oscillations will occur there is also interest in
tailoring the relationship between supply or battery voltages and pacing rate to suit
111
the design of the recharging scheme. As time goes on and the lithium cell’s charge is
drained, the supply voltage decreases. This results in a small but detectable change in
the rate of output pulses as the internal voltage drops in the PUT become a larger
fraction of the supply voltage to the circuit. The relationship between output pulse
rate and supply voltage is used to monitor charge level and recharging rate of the
battery. A steeper relationship helps to overcome various sources of noise in the
measurements and estimations but the pacing rate must stay within a physiologically
desirable range.
There are two main effects on the relationship between pacing rate and supply
voltage, and they are not entirely independent. The first effect is a result of the
relationship between the bias resistors, or their ratio. This is because the ratio sets
up the gate voltage, which determines when the device will fire along the charging of
the output capacitor. If the capacitor reaches the gate voltage near the beginning of
the capacitor charging curve (Figure 5-4) when the curve is quickly rising, the
threshold will be reached more quickly. If the gate voltage is set higher, towards the
end of the capacitor charging curve, then the threshold is reached much less quickly.
The output rate will be lower and also more susceptible to noise in the supply voltage
and the components themselves. The relationship is not linear and works more in
favor of a steeper curve when the gate voltage is set higher. The other effect is choice
of RC. However, RC is not a component that can be changed easily with this design, as
112
it determines rate of the device. Therefore the values of RC that will allow oscillations
will be reviewed (Value of RC and Minimum Current).
Figure 5-4 Relaxation Oscillator Operation
Voltages seen on the relaxation oscillator's output capacitor COUT and at the tissue Ve.
Three different ratios were simulated (open shapes) and then tested on a
breadboard (filled shapes), maintaining similar RG and a similar pacing rate to isolate
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the effect of rate (Figure 5-5). Simulations were done in PSPICE with identical
components. Both simulations and benchtop studies show that it is most favorable to
use the relationship that yields the steepest slope (20%), because this will yield the
best resolution when calculating the battery voltage or charging current.
Figure 5-5 Pacing Rate Vs. Supply Voltage for Different Ratios
Graph of BPM vs. Supply Voltage. A 20% ratio of R1 to R2, or a higher gate voltage, produces a battery
depletion curve with a more drastic slope. This will be helpful when we are using BPM to determine
the current charge level and the charge rate.
The PUT PSPICE model does not operate in the same range of supply voltages as
the physical PUT, so data capture is limited. The PSPICE model seems to have tighter
limits, ceasing function at higher supply voltages. The predicted rate of the device in
40
50
60
70
80
90
100
110
120
0 1 2 3 4 5
Beats Per Minute
Supply Voltage (V)
50%
30%
20%
50%
SPICE
114
PSPICE is usually lower than it is on the physical board. The changes in slope agree
with the prediction that the lowest ratios, and therefore the highest gate biases, are
more affected by the offset voltage of the PUT, and therefore have steeper slopes.
Figure 5-6 Discharge of 3 mAh Battery for Different Ratios
Discharge of a 3mAhr battery over time as seen by pacing rate (BPM). The discharge curve is shown
for three different component sets representing ratios from 20 to 50%.
When this curve is combined with the discharge profile, obtained on the benchtop
with a Quallion Li Ion 3.0mAhr battery, the BPM of the relaxation oscillator can be
predicted over time (Figure 5-6). The discharge profile was taken with a total
discharge current of about 13uA. This graph also shows that it is more favorable to
60
70
80
90
100
110
120
0 50 100 150 200 250 300
Predicted Beats Per Minute
Time (Hours)
BPM 50%
BPM 30%
BPM 20%
115
use a lower ratio because of a more gradual and detectable reduction in supply
voltage.
Value of RC and Minimum Current
Another relationship that affects the choice of component values is the inability to
generate oscillations if the capacitor charging resistor, RC (Figure 10 4), is not large
enough. This low component value creates a situation where the PUT won’t switch to
its high resistance “off” state and the output capacitor Cout cannot recharge.
Therefore the oscillator becomes stable and no longer oscillates. The value of RG
affects this relationship because it defines the valley current IV and therefore R1 and
R2 have an impact on this choice of RC as well. Simulations of working parameters
have been observed as being more conservative than the devices on the benchtop, so
working values were not able to be determined from PSPICE. Therefore, multiple
component combinations were generated with different values of RG and RC, keeping
ratio near about 20%, and the best combination was chosen. The data generated from
this iteration can be observed in Table 5-1 Component Selection Table. The
component values chosen were those with R1 to R2 ratio of about 1:4 and a rate
nearest to 100. These values were changed slightly when fabricating the small
pacemaker device to accommodate the smaller pacemaker layout.
116
R1
(kΩ )
R2
(kΩ )
R1/R2 RG RC
(kΩ )
C
(μ F)
Rate
(BPM)
Pulse
Max (V)
Charge
(μ C )
180 806 0.22 147.14 500 1 60 2.4 2.40
100 360 0.28 78.26 330 1 100 2.4 2.40
100 360 0.28 78.26 300 1 103 2.4 2.40
100 360 0.28 78.26 360 1 92 2.4 2.40
200 820 0.24 160.78 500 1 65 2.44 2.44
100 360 0.28 78.26 360 1 94 2.44 2.44
100 500 0.20 83.33 360 1 80 2.6 2.60
100 500 0.20 83.33 330 1 80 2.6 2.60
100 500 0.20 83.33 307 1 86 2.6 2.60
100 500 0.20 83.33 300 1 86 2.6 2.60
100 500 0.20 83.33 285 1 88 2.6 2.60
68 360 0.19 57.20 330 1 75 2.6 2.6
68 360 0.19 57.20 240 1 103 2.6 2.6
Table 5-1 Component Selection Table
Various combinations of components were analyzed to determine the best combination for the fetal
pacemaker. This was an essential step because the PUT is being used at the edge of its capabilities,
meaning that the data sheet parameters cannot always adequately predict operation. Data generated
by breadboard when testing various component combinations. The design demanded a pacing rate of
around 100 BPM and a R1/R2 ratio near 20%. Therefore, the last row’s parameters were chosen for
the pacemaker design. These were later adjusted to accommodate the smaller pacemaker layout.
Accepted Component Values
The current design of the pacemaker electronics uses a safety factor of over three,
delivering a charge of 2.6 µC per stimulus with the component values found in Figure
3-3. Based on these values, the fully charged lithium cell is calculated to support about
1 to 2 weeks of operation, which is similar to what is seen on the bench during testing.
The values that are currently being used in production may be changed if the charge
117
being delivered is not sufficient to stimulate in all cases as determined by animal
studies. COUT could be made larger (1.5 µF) and RC would be reduced to achieve the
desired rate. If stimulation thresholds are found to be much lower, COUT could be
reduced and RC increased in order to prolong battery life.
Recharging During Fabrication: Benchtop System
Design
The requirements for the recharge system are different between fabrication and
implantation stages of the device. During fabrication, the device is on the benchtop,
allowing strong coupling and direct rate measurement in a simple circuit simulating
the output load. During implantation, the device is much farther from the primary coil,
reducing coupling and adding dissipative body tissue. The pacing rate is affected by
the complex and variable load of the electrodes and tissue (see below). The implant
will be recharged by driving a small diameter primary coil with low power while the
device is on the benchtop and directly probing the output electrodes of the device to
measure rate. Once the device is implanted, another primary coil large enough to fit
around the mother’s abdomen will be driven by a high power Class E oscillator. Rate
in the implanted case will be measured by use of a custom designed electrical impulse
detection amplifier that picks up an artifact from the pacing impulse on the surface of
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the mother’s abdomen. Pacing pulses will be fed into software to calculate rate and a
control algorithm will monitor charge level and recharging rate of the battery.
The initial development and validation of the recharging control system was
conducted with the benchtop recharge system, which avoid some of the complex and
poorly controlled conditions that arise in vivo. This algorithm is depicted in Figure
5-7. A primary coil was fabricated to fit around the device plus its sterile packaging.
The coil is driven at 6.78 MHz, a narrow ISM band allowing unlimited RF power, by
an off-the-shelf function generator (Global Specialties PW2120) with a maximum
power output of 200mW. The pacing rate is calculated in LabView (National
Instruments, 2009). ). Validation in vivo is described in a subsequent chapter.
Testing
Successful charging of a device is shown in Figure 5-8. A LabView program was
used to implement the algorithm. The battery voltage is extrapolated from the rate
information by using the calibration curve (right). Battery voltage over time (left,
white data) shows a charging curve that is typical of lithium ion secondary cells, and
recharge current (left, red) stays around 0.9mA.
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Figure 5-7 Recharging Algorithm
Algorithm for monitoring the pacemaker rate and therefore supply voltage. The flowchart displays a
control in battery voltage and recharging current.
Yes
No
Start
Yes
No
Turn off carrier
IDLE
Monitor
BPM
Decrease the
carrier output
(-1V)
No
Yes
BPM<
104?
Yes
No
Carrier Off
bpm1-
bpm2
>3.2
2.5<=
bpm1-
bpm2
<=3.2
BPM<
108?
Turn on carrier 10sec
Measure bpm1
Turn off carrier 10sec
Measure bpm2
Turn on carrier
60sec
Increase the
carrier output (+1V)
120
Figure 5-8 Benchtop Recharging System Front Panel
(Left) The battery voltage and recharging current is displayed over time for a recharge session. (Right)
An example of the calibration curve generated before monitoring begins.
Next Generation Design: High Power Driver System
The benchtop system works well to keep cells charged prior to implantation.
However, a much more powerful driver is required for use in the clinical system, as
well as a rate detection system. This high power recharging system will be discussed
in Chapter 6.
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Chapter 6: Design and Testing of a Transcutaneous RF Recharging
System
Adriana N. Vest, Li Zhou, Xuechen Huang, Viktoria Norekyan, Glen Griffith, Yaniv Bar-
Cohen, Ramen H. Chmait, and Gerald E. Loeb
Preface
This chapter contains the manuscript that describes the transcutaneous, and
therefore high power inductive system that is intended to be used clinically to
recharge the pacemaker after implantation.
Abstract
We have developed a rechargeable fetal micropacemaker in order to treat severe
fetal bradycardia with comorbid hydrops fetalis. The necessarily small form of the
device, small patient population, and fetal anatomy put unique constraints on the
design of the recharging system. To overcome the aforementioned constraints, a high
power field generator was custom built and the recharging process was controlled by
utilizing pacing rate as a measure of battery state, a possibility because of the
relaxation oscillator used to generate stimuli. The design and verification of the
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recharging system is presented here, showing successful generation of recharging
current in a fetal lamb model.
Introduction
We have developed a fetal micropacemaker to treat progressive complete heart
block with comorbid hydrops in the human fetus. This rare condition only occurs in
several hundred pregnancies a year in the United States, but is life-threatening; once
hydrops develops as a result of heart block, fetal demise is nearly inevitable if the
fetus cannot be delivered due to prematurity or other clinical concerns.[28-30, 38,
54] Implanting our fetal micropacemaker could reverse the course of this condition,
result in the resolution of hydrops within one to two weeks by pacing the heart and
restoring adequate blood flow to the fetus. The fetus could then proceed with an
otherwise normal gestation and delivery. Once born, the infant would be implanted
with a standard adult pacemaker and epicardial lead, the standard of care for
newborns with symptomatic bradycardia.
Several groups have attempted to pace these fetal patients in the past, but have
been unsuccessful.[35-37, 55, 56] The causes of failure and eventual fetal death were
not clearly identified, but likely due to three possible reasons:
• Complications from open surgery [55, 56]
• Lead placement complications [35, 36]
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• Lead dislodgement[35] or strangulation
In order to overcome these failure modes, a successful fetal micropacemaker must
utilize a minimally invasive technique in order to reduce the chance of surgical
complications, and the device must be implanted entirely within the fetal chest,
avoiding the complications associated with a trans-uterine lead.
Figure 6-1 Pacemaker Insertion Strategy
(A.) Insertion cannula and trocar (B.) Pacemaker before deployment in insertion sheath (C.)
Pacemaker after deployment.
Our device overcomes the shortcomings of previous attempts and addresses the
requirements by being implanted via a percutaneous approach through a standard
fetal surgical cannula and with a form factor that can fit entirely in the chest wall
(Figure 6-1). The fetal instrumentation necessitates a cylindrical form that fits
through the inside diameter of the 3.8 mm cannula. The potential market size of 500
124
devices a year further constrains its engineering; the technology must be simple and
inexpensive both to develop and to build.
Figure 6-2 Schematic and SPICE Simulation of Pacemaker
The pacemaker provides an exponential decay pulse utilizing a relaxation oscillator.
We employ a simple relaxation oscillator based on a single transistor (PUT, Figure
6-2), which uses only 7 components and keeps the pulse generator small while also
meeting requirements for low power consumption and low development cost. Output
capacitor Cout is charged through resistor RC until its voltage reaches the threshold
set by the programmable unijunction transistor (PUT) and biasing resistors R1 and
R2. The PUT then switches to a low impedance state that allows the charge
RC
125
accumulated on Cout to discharge through the electrode, represented here as a
capacitor in series with a resistor (Ctissue and Rtissue). Once Cout completely
discharges, the PUT switches back into a high impedance state and begins another
charge/discharge cycle at a period set by the time-constant RC*Cout. The output
pulse decays exponentially with a total charge that is fixed by the electronic
component values (Cout and the bias resistors R1 and R2) and supply voltage, and is
independent of electrode impedance.
The rate of the oscillator is nominally set to about 100 – 150 bpm in order to
provide an adequate heart rate to the distressed fetus, essentially functioning in fixed-
rate pacing mode (i.e., VOO). However, the rate also varies to a lesser degree
according to the supply voltage. As the cell's charge is drained, its voltage decreases,
resulting in a small but detectable change in the rate of output pulses as the internal
voltage drop in the transistor becomes a larger fraction of the supply voltage to the
circuit. This relationship is calibrated during manufacture of each micropacemaker
(Figure 6-3), and used to determine the state of the pacemaker cell, and therefore
when to recharge. This technique was used to decide when to replace the earliest
cardiac pacemakers circa 1960, which were based on a relaxation oscillator circuit
similar to what we have used in this device.
126
Figure 6-3 Calibration Curve
A typical calibration for a pacemaker. Each pacemaker is serialized and then calibrated so that the
pacing rate can be used as a measure of cell voltage.
For power, the oscillator uses a tiny cylindrical battery with a capacity of 3 mAh
(QL0003I Zero-Volt, Quallion), which can sustain a typical stimulus output pulse of 3
µC for 6 days. Therefore, the pacemaker must be transcutaneously recharged
periodically over the device lifetime, which is intended to be from 1 to 3 months.
Electromagnetic induction, a mature wireless power methodology used in biomedical
applications, can be used to transmit energy to the fetal micropacemaker through the
maternal abdomen, recharging the lithium ion cell. This method relies on inductive
coupling between a primary and secondary coil, but is complicated in this application
by the highly variable and randomly changing distance and orientation between them
as a result of fetal movement in the uterus. The shape and size of the secondary coil
127
are severely constrained by the implant form factor and packaging while the size and
orientation of the primary coil are constrained by the anatomy of the mother.
This paper describes the design of the recharging system while addressing the
design constraints that are inherent in a fetal micropacemaker system, and then
details the verification of that system.
Design
Electromagnetic Field Generation
Because the coupling is expected to be very poor due to a large distance between
coils and a very small secondary coil, the implant needs the largest permeability,
diameter, number of turns and operating frequency that it can attain in order to
generate as much electromotive force (EMF) as possible. The diameter of the
secondary coil (2.75 mm) is already constrained by the required dimensions of the
device. By simultaneously solving for Q and induced EMF, 23 turns of 3mil wire was
found to generate sufficient current in the implant while not limiting the bandwidth
of the device. An operating frequency of 6.78MHz was chosen because it is high
enough to allow an appreciable amount of inductance without too many turns, is low
enough to be efficiently generated with modern timing circuits, and is at an ISM band,
128
allowing unlimited radiated power
1
. To further increase permeability and therefore
induced EMF, the secondary coil is wound around a ferrite core made of material 61.
Through several design iterations we found that maximizing the length of the ferrite
core within the package constraints further improved coupling.
To multiply the voltage across the load (the battery), a parallel resonant circuit
topology is used (Figure 2, Ltank and Ctank); a resonator Q of about 4-5 allows for
operation without precise tuning and yet provides the sought after voltage
amplification. A half wave rectifier (D1) maintains a higher equivalent RMS
impedance so that the inductance of the implant coil can be allowed to be larger and
maintain the Q, while also providing a minimum component count. The RMS
equivalent load is half of the DC resistance, neglecting the rectifier diode drop voltage.
The lithium cell would require a 4.2 VDC output from the coupling elements and 1.5
mA of current (C/2 recharging speed); this results in a DC resistance of 2800 Ω, and
an AC RMS load of 1400 Ω. An inductor in parallel configuration would need a
reactance of 350 Ohms to maintain a Q=5.5. This corresponds to an inductance of
8.22uH. The necessary open circuit EMF of the implant coil would be 4.2V/5.5 = 0.76
V.
1
FCC guidelines allow unlimited radiated power at ISM bands, but limit power at other bands in
order to prevent EM interference. Other FCC guidelines exist that limit human exposure to EM fields,
and these guidelines also apply to ISM bands. Therefore, analysis will need to be performed to ensure
the field strength does not exceed this limit.
129
To determine the required electromagnetic field to generate the necessary EMF,
we use Faraday’s Law of Induction. The changing magnetic field induces an
electromotive force, ε, in the implant according to
𝜀 = −𝑁 2
µ
𝑒
𝑑 𝑑𝑡
(𝐵 ∙ 𝑆 ) Volts
where N2 is the number of turns of the secondary coil, µe is the effective permeability
multiplier, B is the magnetic field density, and S is the surface area of the implant
through which the flux is passing.
𝑆 = 𝜋 ∗ 𝑟 2
meters
where r is the radius of the secondary coil. By the Biot-Savart law, and a simplification
of the magnetic field on axis with a number N1 of current loops on the primary
𝐵 = 𝑁 1
∗
𝜇 𝑜 𝑅 2
𝐼 sin (𝜔𝑡 )
2 (𝑧 2
+𝑅 2
)
3
2
Teslas
where µo is the permeability of free space, R is the radius of the primary coil, I is the
current in the primary, ω is the radian frequency, and z is the distance from the
primary coil on axis. Because the current, and therefore the magnetic field, is
sinusoidal, the time derivative yields the following:
𝜀 = −𝑁 2
𝜇 𝑒 ∙ 𝜔 ∙ 𝐵 ∙ 𝜋 ∙ 𝑟 2
Volts
The frequency, permeability, and secondary coil are constrained at this point in
the design. Therefore, in order to generate the required EMF at the required distance
of about 10 cm, we can vary current, number of turns in the primary coil, and
diameter of the primary coil. It is desirable to keep the number of turns in the primary
low to minimize dissipative and dielectric losses in the coil turns. An external coil
130
radius of 20cm yields the best balance of maximum field range combined with
minimal coil power dissipation and coil voltage (Figure 6-4).
Figure 6-4 EMF Vs. Distance from Primary Coil
(Left) Radius (R) is varied and (Right) current (I) is varied of the primary coil to determine the EMF
that can be generated in the implant.
To generate the oscillating current necessary as found above, a Class E Oscillator
was selected for its highly efficient operation and tolerance for operating point shifts.
The field generator uses a Pierce Gate Oscillator and a 6.78MHz crystal to generate
the correct frequency for the ISM band in which it is desired to operate. The
oscillator’s signal is buffered through a network to generate enough current to drive
the N-Channel MOSFET in the Class E Oscillator. The oscillator uses resonance to
further increase the power of the signal. The MOSFET is configured as a common
source amplifier and current is drawn into the drain from a power source through a
5.6 µH RF choke. The power supply voltage can be varied, resulting in the generation
of different strength fields, a feature that would become useful if the pacemaker was
being recharged too quickly and the field was too strong. This situation would arise if
131
the pacemaker was oriented very favorably and was closer in distance to the primary
coil. The remainder of the Class E amplifier includes capacitors and an inductor
required to generate the required resonant frequency and produce the equivalent of
a 50 Ω resistive source impedance for efficient connection to the transmitting loop
antenna via coaxial cable.
Figure 6-5 Schematic of the Coil Driver and Antenna
A Class E Oscillator (left) generates the 6.78MHz sine wave that drives the Antenna (left).
The transmitting antenna uses a balun (L1, L2) to take the unbalanced output of
the Class E Oscillator and create a balanced input to the center tapped resonant coil,
effectively dividing the peak voltage across the coil, which adds to its safety. A
network of high voltage RF capacitors provides tuning to resonate the antenna coil
and to create a 50 Ω input impedance to match the above mentioned source. The coil
is realized by two turns of coaxial cable, thereby providing shielding to reduce
radiation of electrostatic fields.
132
Recharging Control Scheme
Due to the small size of the implant and limited resources for regulatory circuitry,
the recharging system must be regulated without adding too many components.
Therefore, the strategy employed uses only external monitoring to regulate the
recharging process. The pacing rate provides a measure of the voltage supplied by the
pacemaker’s Li Ion cell at any time after calibration, which can be used to determine
the state of charge of the cell from the discharge profile. The changes in output rate
are subtle and actual rates depend on component values that are somewhat variable,
so a precise calibration curve of output rate vs. supply voltage is created during
manufacture of each device (Figure 6-6).
Figure 6-6 Diagram of the Recharging Strategy
The Artifact Detection System monitors the pacing rate of the pacemaker and infers battery voltage,
given the serial number of the device. A Modulatable Field Generator can be adjusted if the recharging
is determined to be too fast or too slow.
133
Rate is measured in vitro, during fabrication and testing, by a simple benchtop
system and LabView software that clocks the interval between each stimulus. Rate is
measured in vivo by sensing the conducted pacing artifact via skin electrodes placed
on the maternal abdomen. The signals are acquired and transmitted to a PC by the
BioRadio Wireless Physiological Monitor (Great Lakes Neurotechnologies) and
analyzed with the same LabView software used for the in vitro measurements (Figure
6-6).
Using cell voltage and the datasheet, the state of charge can be inferred and is used
to determine when the pacemaker needs to be recharged and when the pacemaker is
fully charged. It is also important to limit the maximum charging current to prevent
damaging the lithium cell. The current varies depending on the strength of coupling
between the primary and secondary RF coils. Because of the limitations of the field,
recharge current is typically less than the maximum specified by the manufacturer.
In a more powerful design iteration of the field generator or with a superb
pacemaker position current may exceed the datasheet limits. In this case, it is useful
to measure charging current. This can be done in vivo by observing potential
developed across any resistance in series with the cell according to E = IR. A series
resistor RB (Figure 6-2) was added to the circuit for this reason. For example, a 3 V
lithium cell with a combined internal resistance and series resistance of 1 kΩ that is
134
being recharged with 1 mA of current will actually generate 4 V across the relaxation
oscillator, causing the oscillator to cycle more quickly. By accurately measuring the
output pulse rate when the recharging field is turned on and again when it is turned
off, this voltage change and the actual charging current can be computed. If that
recharge current is excessive for the ratings of the cell, then the field strength
produced by the external transmitter can be reduced manually or programmatically
until the rate change is consistent with the desired recharging current. If the
recharging current is quite low, the operator can be advised to try to reposition the
primary coil on the abdomen.
The charging current calculation can sometimes be unreliable because the intense
EM field also causes some changes in rate simply by adding noise to the electronic
circuit, confounding the relationship between cell voltage and pacing rate. Therefore,
another method is used to calculate current, by measuring the change in battery state
of charge over time; State of charge is assessed, the EM field and recharging current
is applied for a set amount of time and then turned off, and state of charge is measured
once again. The difference in state of charge is divided by the amount of time of the
recharge, and current is calculated. Sudden changes in pacing rate during charging
that are caused by fetal motion can still be detected and used to trigger corrective
actions.
135
Calibration
The pacemaker must be meticulously calibrated in order to keep track of the
relationship between the supply voltage and the pacing rate. Each device is calibrated
initially by varying an external DC supply voltage and measuring the pacing rate at
room temperature with a simulated tissue load, a 1 kΩ resistor. Several factors affect
this calibration in addition to cell voltage (VB), including the temperature of the
device, output impedance, light, electrical noise (especially from the high frequency
recharging coil), and recharging current.
To simplify consideration of the factors, we consider here only temperature and
output impedance. The factors of light, electrical noise, and current can be eliminated;
the pacemaker can be shielded from light and noise during calibration, and current is
not present without the recharging field. Temperature and output impedance are the
main factors that change immediately upon implantation of the device, because the
body tissue is elevated in temperature to about 37
o
C and the electrode-tissue
interface is comprised of both resistive and capacitive elements. Both temperature
and the addition of a capacitive element raise the output rate with respect to supply
voltage. The capacitance of the electrodes (the myocardial pacing electrode and
reference electrode provided by the battery case) adds in series to the output
capacitance of the relaxation oscillator, causing the oscillator to see a reduction in
capacitance and therefore a speeding up of pacing rate.
136
Methods
To verify that the recharging system design met requirements for accurately
determining supply voltage and effectively recharging the Li ion cell, a combination
of in vitro and in vivo tests were completed. These tests measured the field generated
with the recharging system and the current generated in the implant in vitro, assessed
the environmental effects on calibration in vitro and in vivo, and evaluated the
performance of the recharging system in vivo.
The magnetic field generated with the recharging system was mapped by
measuring the signal coupled into a custom made loop probe fashioned out of two
turns of AWG 15 magnet wire and calibrated based on a known magnetic field. A 0.9%
saline bath was used as a tissue model to simulate the conduction of electrical signals
through the body and the loading effect that the body has on the primary coil. To
determine the current induced into the lithium ion cell the field strength was held
constant and an ammeter (Agilent 34410A) was added to the pacemaker circuit
between elements RB and the cell. A twisted pair was used in order to minimize the
effect of the added circuit element on the system in the large electromagnetic field.
To assess the effects of temperature and impedance on the pacemaker calibration
in vitro, a pacemaker was subjected to various environments and pacing rate was
measured. The in vitro conditions included the combination of a simulated tissue load,
either a 1 kΩ resistor or 0.9% phosphate buffered saline (PBS), and an oven
137
controlled temperature environment, either room temperature (23oC) or body
temperature (37-40oC). Pacing artifact was observed across the simulated tissue load
using an amplification circuit or the biopotential recording system (BioRadio 150,
Great Lakes NeuroTechnologies) with silver chloride electrodes placed at the surface
of the saline. The artifact signal captured was digitized, captured on PC, filtered to
eliminate DC drift and high frequencies (LPF=400 Hz, HPF=5 Hz), and the period of
the stimulus artifact was measured.
The effect of temperature and impedance on calibration was assessed in vivo by
observing pacing artifacts over several days during animal studies. The in vivo study
protocol was approved by the Institutional Animal Care and Use Committees at the
University of Southern California and the Los Angeles Biomedical Research Institute.
Fetal micropacemakers with rates between 100 and 150 bpm were implanted into
seven fetal sheep according to the procedure discussed in Bar-Cohen et al. [57].
Devices were implanted fully charged, but supply voltage and therefore pacing rate
decreased over time, providing observations of pacing rate at different voltages.
Because the fetus was instrumented for these experiments, the fetal micropacemaker
artifact was detected by three trans-uterine electrodes attached directly to the fetus.
The percutaneous leads extended out of a maternal abdominal incision and into a
pouch sewn on the maternal skin surface for interfacing with the above mentioned
biopotential recording system for analysis.
138
The performance of the recharging in vivo was assessed by quantifying the shifts
in pulse rate described above. After implantation of the pacemaker and during follow
up days, recharging was attempted at the discretion of the animal follow up team.
Ultrasound imaging was performed to determine the location and orientation of the
pacemaker. The primary coil was placed on or around the maternal sheep to minimize
the distance and angle between the primary and secondary coil. Increases in pacing
rate during field generation provided a measure of the coupling between the two coils,
and was used to position the primary coil. The value of resistor RB was increased in
the latter implants (Sheep 5 through 7) to improve this measure.
Results
Magnetic Field and Current Generation
The magnetic field generated by the hardware detailed above as detected with the
loop probe matched well with simulations (Figure 6-7). The near field reached the 30
A/m that the system was designed for and the decay of the field was commensurate
with the simulated decay over distance.
139
Figure 6-7 Field Measurements In Vitro
(A) The electromagnetic field dropped off with distance from the coil (on axis) and increased with
distance from the center (radial). (B) The measured field strength at the center of the primary coil and
as distance increased from the plane of the coil (on axis) matched the theoretical calculation.
In early design iterations, in vitro measured current was 1.15 mA in a field of 30
A/m. The ferrite core was lengthened and current limiting resistor RB was reduced
to improve current induced in the implant. In the final design iteration, in vitro
current measured 2 mA in a 30 A/m field. Fig 7 B shows current measured coaxially
with the coil as distance from the coil increases.
Calibration
The pacemaker used in sheep study 7 was analyzed in vitro and demonstrated
variations in pacing rate due to temperature and impedance changes. The pacing rate
increased with a nearly linear offset by an average of 3.4 bpm from temperature alone,
0
10
20
30
40
On Axis Distance
(From plane of coil)
(cm)
Field Strength (A/m)
Radial Distance
(From Center of Coil)
(cm)
5
10
15
20
25
30
35
0 10 20
Field Strength (A/m)
On Axis Distance
(From plane of coil) (cm)
Measured
Theoretical
Radial Distance = 0
140
9.2 bpm from phosphate buffered saline alone, and 11.0 bpm from the combined
effects of temperature and saline together (Figure 6-8).
Figure 6-8 Effects on Calibration
A pacemaker originally calibrated (red squares) is subjected to various environments to determine the
effects on the calibration curve. The effects of the body (yellow filled in circles) cannot be completely
simulated in a warm saline bath (blue stars).
When the pacemaker was implanted into sheep 7 on operational day (OD), an
initial offset of 10.6 bpm was observed, similar to the offset observed in vitro in 39 oC
PBS. On post operational day 2 (POD2) the pacing rate was observed to have a new
offset of 12.9 bpm. The measured increase of pacing rate offset from OD to POD2
indicates that the effective capacitance of the oscillator circuit was decreasing. This
means that the capacitance at the frequencies of interest of the electrode-tissue
interface was decreasing, perhaps related to the change in the electrode-tissue
90
95
100
105
110
115
3 3.5 4 4.5
Pacing Rate (BPM)
Battery Voltage (V)
1kOhm
Equation for
Calibration
1kOhm, 39oC
23oC, PBS
39oC, PBS
Warm Tissue
141
interface as various biochemical and cellular components of the foreign body reaction
were deposited. This is consistent with data in the literature that show increasing
impedances at low frequencies, a reflection of decreasing interface capacitance [58].
Quantifying the exact changes to the impedance of the system is out of the scope of
this research. However, based on our observations in vivo, these changes seem to
stabilize within 2 days after implant sufficiently to assume an overall capacitance
change, and to select an offset value for the calibration.
Four different implanted pacemakers generated data to enable measuring the
linear offset over time, one in sheep 4, two in sheep 5, and one in sheep 7. The original
calibration curve (dotted line, Figure 6-9) was fitted with a two term exponential
function and then linearly shifted by a factor proportional to the relative pacing rate,
as each device has a slightly different rate based on actual component values. The
offset calibration curve (solid line, Figure 6-9) predicted with good results what the
pacing rate would be on follow up days after OD (discrete data points, o).
142
Figure 6-9 Corrected Calibration Curves
Calibrations (dotted lines) for different devices were linearly shifted (solid lines) to predict a new
calibration for an in vivo environment. Data points taken during in vivo testing are plotted (circles),
indicating the quality of the prediction. Full effect of the in vivo scenario was not complete on the first
day (operational day, OD), hence the offset of the actual value from the prediction.
In Vivo Recharging
Recharging efforts in sheep 4 were not successful and did not result in any
observed change in state of the battery. It was later determined that the recharging
system hardware had some tuning and heating malfunctions, which were solved by
replacement of components, retuning, and long duration burn-in tests.
143
In sheep 5, recharging efforts on POD12 were successful at waking up the implants
after they had been completely depleted. The two devices were nearly side by side,
and recharged to about the same level. Device I recharged to 2.6V and Device II
recharged to 2.8V. The device implanted into sheep 6 was able to be recharged to
between 3.0 to 3.4V on POD2. Recharging was performed for just a short time in sheep
7, but changes in pacing rate were observed. Low speeds of charging (current) were
observed in all recharging attempts.
Discussion
The recharging system demonstrated here successfully generated the expected
electromagnetic field. Realized currents, even in the latest design iteration, were
slightly lower than calculations predicted. This is due to simplifying assumptions
made in the initial analysis for coupling and Q, and could account for the low currents
seen during in vivo experiments as well. The initial design considerations assumed
an implant Q of 5.5. However, this value actually varies with the conduction of the
diode. When the diode is conducting, a very high current would actually load the
secondary coil heavily, bringing down the Q of the device. The value of RB also limits
the amount of current generated into the implant and was reduced in devices built
after animal studies to increase current, at the expense of a less sensitive indicator of
recharge current. Increasing field strength would improve current generated in the
implant and could be done by increasing the number of turns on the primary coil,
144
increasing current on the primary, and employing components with higher power
ratings.
The linear offset of the calibration to accommodate environmental effects fits the
data on the limited implants tested for the limited durations tested. Further in vivo
tests to ensure the proposed calibration shifts are stable would be useful for
validation of this approach. Future work will involve incorporating environmental
factors that were not considered here, such as light, electrical noise, and a greater
sample of temperatures and impedances, leading to the formation of a multivariable
calibration equation.
Results from the in vivo recharging efforts were promising, although hardware
malfunctions and short subject follow up periods led to limited data collection.
Current in devices was induced in vivo, as evidenced by the change in voltage over
time of the pacemakers. However, this current was less than expected. Changes in
device design to amplify the induced current were completed after animal studies had
concluded, so the in vivo studies were not able to take advantage of these
improvements. Anatomy of the pregnant ewe and poor follow up imaging quality
made positioning the primary coil near the implant difficult, further exacerbating
coupling problems. Time was limited during follow up visits because of animal
handling restrictions, and the low currents could not be compensated for with longer
recharge sessions.
145
These challenges will be much more easily surmountable in a human scenario
which will use the latest pacemaker design iteration because imaging quality will be
much clearer, the pregnant mother will be responsive to direction, and time with the
patient recharging is able to be extended in the face of low induced current. Future
work will include repeated in vivo recharging of devices over time to validate the
overall system as well as the proposed calibration factor.
Acknowledgments
The authors would like to thank Drs. Michael Silka, Jay Pruetz, and Catalina Guerra
for assistance in imaging, implanting devices, and clinical expertise; consultant Ray
Peck for fabrication and engineering expertise; and engineers Kaihui Zheng, Michael
Lu, Hithesh Reddivari, Shane Garcia, and Michael Maylahn for software and electronic
design and animal study preparation contributions. This work was generously funded
by the Coulter Foundation, the NIH (grant #: 5R01HD075135-02), the Wright
Foundation, and Southern California Clinical and Translational Science Initiative.
146
Chapter 7: Preclinical Testing and Optimization of a Novel Fetal
Micropacemaker
Yaniv Bar-Cohen, Gerald E. Loeb, Jay D. Pruetz, Michael J. Silka, Catalina Guerra,
Adriana N. Vest, Li Zhou, Ramen H. Chmait
© 2015 Elsevier. Reprinted from:
Bar-Cohen Y, Loeb GE, Pruetz JD, Silka MJ, Guerra C, Vest AN, Zhou L, Chmait
RH. Preclinical Testing and Optimization of a Novel Fetal Micropacemaker.
Heart Rhythm. In Press.
Preface
This chapter covers the methods and results from preclinical testing of the fetal
micropacemaker in a fetal sheep model. During these experiments, the mechanical
and electronic aspects of the device were tested. To ensure successful pacing of the
fetus was accomplished, fetal ECG was measured. Additional electronic results about
the specifics of the recharging system tests can be found in Chapter 6 and additional
physiologic assessments can be found in Chapter 8.
Glossary of Abbreviations
CHB: Complete Heart Block
147
Abstract
Background: Fetal complete heart block with hydrops fetalis is a life-threatening
condition with no effective treatment options when early delivery is not feasible. We
have developed a novel rechargeable fetal micropacemaker that can be implanted
percutaneously into the fetus.
Objective: To implant and follow-up the micropacemaker device in a fetal sheep
model.
Methods: Percutaneous pacemaker implantations were performed in seven fetal
sheep (112-128 days gestation). Animals with successful implantations were
followed until death or anticipated delivery. During follow-up, external battery
recharging via a wireless system was also trialed. In the last four animals, a
pericardial effusion was created via needle-injection of saline prior to implant in
order to improve imaging and device placement.
Results: Despite limitations of the fetal sheep model, successful ventricular
capture was demonstrated in five of seven fetuses. Increasing capture thresholds
were demonstrated in the follow-up period (5 to 19 days), and subsequent histology
was notable for a marked inflammatory response at the site of electrode implant.
Wireless battery recharging was attempted in the final four animals, and after
optimization of the system, recharging was demonstrated in the last three. While
premature delivery / death occurred in three animals during the follow-up period,
there was no evidence that these events resulted from the percutaneous pacemaker
or its implantation.
148
Conclusion: Our results demonstrate the potential promise of a completely
implantable, percutaneous micropacemaker for the treatment of complete heart
block in the fetus. Fetal stimulation thresholds and subsequent tissue responses
appear similar to post-natal parameters.
Introduction
Complete heart block (CHB) in the human fetus is a life-threatening emergency
with no effective treatment options beyond watchful waiting. [30] Fetal bradycardia
due to heart block may be progressive in utero, and hydrops fetalis may develop in
more than a quarter of these pregnancies. [40] Once hydrops fetalis occurs, if the fetus
cannot be delivered due to prematurity or other clinical concerns, fetal demise is
nearly inevitable. Pharmacologic and immunologic therapies (including
corticosteroids, beta agonists and intravenous immunoglobulins) have been
undertaken in an effort to treat these fetuses with limited, if any, positive effects or
survival improvement. [33, 40, 59]
Although the degree of myocardial dysfunction that can occur due to antibody
damage in congenital CHB cannot be completely predicted, successful pacing of a
fetus with CHB and hydrops fetalis could theoretically allow resolution of hydrops in
several weeks and permit an otherwise normal gestation. Historical attempts at
pacing human fetuses, however, have invariably failed with no survivors to date. [34-
149
37] We have designed a single-chamber pacing system (Figure 7-1) that is self-
contained and can be percutaneously implanted in the fetus without exteriorized
leads, thereby permitting subsequent fetal movement without risk of electrode
dislodgement.
Figure 7-1 Fetal Implantation Equipment
Photographs of the implantation equipment. A: Implantation cannula with trocar inside. B: Sharp tip
of trocar protruding from end of cannula. C: Micropacemaker inside its implantation sheath
protruding through cannula. D: Micropacemaker device with distal electrode screw connected to the
micropacemaker (3.475 mm diameter, 18 mm long) via a coiled flexible lead.
150
Previous preclinical animal studies in adult rabbits have demonstrated the
viability of our implantation scheme and have allowed us to optimize the electronics
of our pacing system.[60]
We now report on implantation outcomes of functional
micropacemakers in fetal sheep.
Materials and Methods
Device Design and Function
In order to meet our rigorous requirements regarding device size, power
consumption and development cost, we utilized a simple relaxation oscillator based
on a single transistor.[60] The device functions in a fixed-rate mode (i.e., VOO) with a
rate that predictably varies with the battery voltage (generally 100-110 bpm). The
output pulse exponentially decays with a total charge that is fixed by the electronic
component values and is independent of electrode impedance. In 6 of 7 experiments,
the devices had a 3 µC output pulse (3V peak, 250 µs time constant); this corresponds
approximately to a conventional 2 V over 0.4 ms square pulse. The 3mAh lithium
battery cell in the implants sustain pacing for approximately 6 days and can be
recharged by inductive coupling of a 6.78 MHz electromagnetic field from a
transmitting coil positioned outside the maternal body. With the 6th implant, a higher
output device (7.2 µC; 3-day recharging interval) was used.
151
Device Implantation
Implantation experiments were performed on seven pregnant fetal sheep (ovis
aries, Rambouillet and Columbia mix breed) at 112-128 days of gestation with only
singleton pregnancies included. The protocol conformed to the Guide for the Care
and Use of Laboratory Animals and was approved by the Institutional Animal Care
and Use Committees at the University of Southern California and the Los Angeles
Biomedical Research Institute. The ewes were anesthetized with ketamine and
atropine, followed by isoflurane. Oxacillin and gentamicin were given as preoperative
antibiotics. Ultrasound imaging was performed on each maternal sheep prior to the
initial incision to assess the fetal position and heart.
Dissection down to the uterine wall was made via a midline abdominal incision,
and the uterus was externalized. A purse-string suture was placed and a uterine
incision was made. The fetal thorax was exposed and electrodes were placed in
triangulated locations surrounding the fetal heart to allow fetal ECG evaluation.
In the first three animals, the fetal subxyphoid area was exposed via the uterine
incision for insertion of the percutaneous implantation trocar and cannula (3.8 mm
internal diameter, 4.5 mm external diameter; Figure 7-1) (Richard Wolf Inc., Vernon
Hills, IL). In the latter four animals, the trocar and cannula were advanced
percutaneously through an intact area of the uterine wall and into the fetus at the
subxyphoid region. Under ultrasound guidance, the trocar and cannula were
152
advanced through the fetal diaphragm and towards the ventricles. Once ultrasound
imaging suggested that the cannula and trocar were against the epicardial surface,
the trocar was removed and the pacemaker implantation system was advanced
through the cannula. After the distal electrode screw reached the epicardium, the
device was rotated clockwise into the myocardium. The device was then released
(Figure 7-2) by slowly withdrawing the cannula while keeping the device in place
with a pushrod. In cases where successful pacing could not be demonstrated with the
initial device, a second and once a third device implantation attempt was performed
in the same animal.
Figure 7-2 Ultrasound: Micropacemaker deployment
The implanted electrode (thick arrow) and implantation cannula (thin arrow) just after deployment of
the micropacemaker.
153
Figure 7-3 Ultrasound: Creating a Pericardial Effusion
A long needle (arrow) is inserted into the pericardial space under ultrasound guidance (A). Saline is
injected in order to create a transient pericardial effusion (B).
After initial experiments demonstrated difficulty in correctly positioning the
implantation cannula at the epicardial surface, a pericardial effusion was created in
the four latter experiments. This was performed in order to more closely replicate
the clinical anatomy in a human fetus with hydrops fetalis. In addition, the presence
of a pericardial effusion facilitated ultrasound imaging for improved cannula
positioning and more accurate device implantation. After the uterine wall was
154
exposed and fetal electrodes were placed, a 17 cm, 21-gauge needle was advanced to
the epicardial surface (Figure 7-3, A). When the needle appeared to have crossed
through the pericardium, saline was injected until ultrasound imaging demonstrated
the presence of a small pericardial effusion (Figure 7-3, B). The cannula and trocar
was then inserted as described above with improved visualization of the epicardial
surface.
After device deployment, the purse string uterine incision was closed and the
uterus was returned to the abdominal cavity. The fetal electrodes were delivered
through a small incision in the maternal abdomen for postoperative access. Final
ultrasound imaging of the fetus was performed to evaluate the heart rate and rhythm,
cardiac function, electrode and device position and overall fetal wellbeing. The sheep
was then awakened from anesthesia, administered antibiotics (Penicillin G), and
allowed to return to the animal housing.
Fetal Sheep Follow-Up and Device Recharging
Repeat ultrasound evaluations were performed in the first 24-48 hours of follow-
up and once to twice per week thereafter. In addition, the externalized fetal skin
electrodes were used to obtain a fetal ECG during follow-up. In the last four animals,
recharging of the fetal pacemaker was attempted using an external recharging ring
that was specifically developed for the micropacemaker.[61] Device position and
155
orientation were evaluated by ultrasound to determine the optimal orientation of the
recharging ring in order to recharge the micropacemaker battery. Due to difficulty in
assessing device configuration by ultrasound, fluoroscopy was used in sheep #6 to
determine the device orientation while the sheep was standing.
The design of the pacing circuit was modified to increase the pacing rate modestly
(5-10 bpm) in proportion to the received strength of the RF magnetic field. By
evaluating the pacing rates when the recharging ring was deactivated (i.e., no external
voltage applied), the battery voltage and state of charge could be determined. Follow-
up and recharging were continued on functional devices until the procedure was
electively terminated or until unplanned delivery occurred.
Evaluation of Pacing
Because the pacing rates in VOO mode (100-110 bpm) were generally lower than
the fetal heart rates (130-180 bpm), ventricular capture was detected by the presence
of an irregular fetal heart rate. This was identified by a pacing stimuli occurring after
the ventricular refractory period and resulting in a “premature ventricular
depolarization.” Fetal heart rate irregularity from ventricular capture could be seen
by ultrasound, but confirmation of ventricular capture was achieved by observing a
pacing artifact at the start of a premature QRS complex on the fetal ECG (Figure 7-4).
The pacing system did not allow for wireless threshold and impedance
156
measurements after implant, but direct measurements of pacing lead capture
threshold and impedances were performed in sheep #5 just prior to euthanasia.
Figure 7-4 Fetal ECG Showing Capture
Demonstration of ventricular capture: Fetal electrocardiogram demonstrating ventricular capture.
Stimulus artifacts from the pacemaker are seen to march out at the pacing rate of 118 BPM while the
intrinsic fetal heart rate is 170 BPM. Advancement of the QRS (arrow) is seen immediately after a
stimulus artifact when ventricular capture occurs.
Device Explantation
The animals were euthanized (pentobarbital / phenytoin solution) after the
follow-up period was completed, and a fetal thoracotomy was performed to
determine the device location (micropacemaker, flexible lead and electrode). In
addition, injury and inflammation of the surrounding tissues were evaluated, and
157
necropsy specimens were sent for histological evaluation. Due to the presence of the
exteriorized fetal electrodes that were sutured directly into the fetus, spontaneous
vaginal delivery was not survivable. When fetal delivery did occur, a fetal
thoracotomy was performed as described above and histology specimens were sent.
Results
Evolution of Acute Implantation Technique
Device implantations were performed in seven pregnant sheep. Table 7-1
summarizes the procedure results, with ventricular capture after device placement
demonstrated in five of the seven animals. With experience gained from each unique
procedure, the pacing system and implantation methods evolved to maximize the
chance for success for the subsequent sheep.
158
Animal
number
Gestation
al age at
implant
(days)
Device
follow-up
time
(days)
Total
procedure
time
Open
uterus
time
# of
devices
Successful
Capture?
Implantati
on
comments
Follow-up
1 112 19 2:15 1:49 1 Yes Failure of electronics
causing intermittent
pacer output.
POD1: Effusions
(pericardial and
pleural) seen on.
POD12: Complete
resolution of effusions.
POD19: Euthanasia.
2 127 N/A 2:50 2:24 3 No Electrode screws did
not penetrate
epicardial surface.
N/A
3 128 N/A 1:40 1:23 2 No First device did not
contact epicardial
surface. Second
device placement
resulted in
myocardial
perforation and
cardiac tamponade
N/A
4 127 5 2:25 0:45 1 Yes Pericardial effusion
created (40 cc)
POD1 and POD3:
Capture and good fetal
health.
POD5: Animal found
delivered and
deceased.
5 128 15 2:30 0:23 2 Yes Pericardial effusion
created (40 cc).
Capture seen on first
device after implant
and capture of
second device
POD1: Good fetal
health, 1 device
captures
POD7: Good fetal
health, no capture
POD12: 2 devices
recharged, no capture.
POD15: Elective
termination
6 120 5 0:50 0:20 1 Yes Pericardial effusion
created (40 cc)
POD2: Intermittent
capture, good health
POD5: Fetal death,
sheep in labor
7 125 6 1:30 0:17 1 Yes Pericardial effusion
created (50 cc)
POD2: No capture,
good fetal health
POD6: Fetal death,
sheep in labor
Table 7-1 Summary of Implants
Seven fetal sheep were implanted with devices, with various outcomes.
159
Due to intermittent failure of device output with the first implant, the circuit board
design was altered and no further electronic device failures were seen. In addition,
there was prolonged exteriorization of a large portion of the fetus outside the uterus
during implant, associated with spontaneous pleural and pericardial effusions and
ascites on post-operative day (POD) #1 which resolved during follow-up.
Exteriorization of the fetus from the uterus was minimized for future procedures, and
significant effusions were not seen in follow-up after the first experiment.
Blunted screw electrode tips were identified as the cause of failure to penetrate
the myocardium with the second procedure and felt to be due to the storage process.
The recharging storage system was altered to avoid tip blunting due to electrode tip
chafing. In addition, refinements of electrode design and screw bevel were made after
a series of experiments on cadaver chicken hearts.
After the third procedure, a pericardial effusion was created for subsequent
implants due to the need to improve imaging and to enlarge the pericardial space for
percutaneous access. There were no major procedural changes in the final four
experiments.
160
Follow-up and Recharging
Five of the seven animals were followed after the acute implant, with two
euthanized at the initial implant (Table 7-1). An increase in capture thresholds during
follow-up was suggested in three animals (Sheep #5, 6, 7) with either intermittent or
no capture seen despite pacemaker stimuli. Recharging was attempted in the last
four animals and confirmed in three. Similar to acute implant, observations during
follow-up resulted in modifications of the system and recharging techniques.
For sheep #4, device recharging was attempted on POD #1 and POD #3. Only a
weak recharging field was seen by the device, due to an incorrect determination of
the micropacemaker orientation by ultrasound imaging in addition to malfunction of
the recharging system. The external recharging system electronics were augmented,
and the pacing circuit was modified to generate a larger increase in the pacing rate in
proportion to the received strength of the RF magnetic field. Device recharging was
demonstrated in subsequent animals.
In sheep #5, lead characteristics were manually determined for the two implanted
devices just prior to euthanasia on POD#15. For the first device, the ventricular
capture was 3.75 V @0.5 ms and 5.0 V @0.2 ms; the impedance was 258 Ohms. For
the second device, the ventricular capture was 2.5 V @0.5 ms and 5 V @0.1 ms with
impedance of 230 Ohms. A higher output device (7.2 µC) was used for the sixth
experiment. For the seventh experiment, a standard output device was used (3 µC),
161
but the electronics modified to allow for higher outputs under the influence of the
recharging field (therefore allowing transient augmentation of pacing output)
Necropsy and Histology
Necropsy was performed at the conclusion of all experiments to determine injury
to the mother and fetus as well as to assess the location and integrity of the
micropacemaker system. Maternal uterine evaluation suggested normal healing at
the purse string incision and only a small scar at sites where the cannula directly
penetrated the uterine wall. Similarly, external evaluation of the fetal chest and
abdomen showed a small scar at the site of cannula insertions. In many cases, the
micropacemaker devices were found to be crossing the diaphragm (Figure 7-5). In
others, the devices lay inferior to the diaphragmatic surface with only the flexible
leads penetrating the diaphragmatic muscle.
162
Figure 7-5 Necropsy: Pacemaker Crossing Diaphragm
Necropsy demonstrating the micropacemaker device (large arrow) lying through the diaphragm. The
electrode is seen penetrating the left ventricular epicardium (small arrow) (sheep #4)
Necropsy identified the ventricular locations where each electrode had been
implanted. By verifying the location of the circular disk at the proximal end of the
electrode screw in relation to the pericardium, determination of whether the
implantation cannula had penetrated the pericardium could be made. In the first and
second experiments, these proximal disks were seen to be entirely within the
pericardium, suggesting that the sheath had penetrated the pericardium. In the
remaining experiments, the disk was seen outside the pericardium.
163
Histology was performed to identify the degree of myocardial inflammation
where the electrodes were implanted. When the electrode had penetrated the
epicardium, a significant inflammatory response was identified (Figure 7-6). In the
fourth sheep where fetal death had occurred, the placenta had diffuse edema of the
chorioallantoic stroma as well as bands of neutrophils and macrophages. The
necropsy report concluded that a placental infection was the most likely etiology of
the fetal demise, although the specific microorganism was not identified. The
necropsies did not identify a clear cause of death for the sheep #6 and 7.
Figure 7-6 Histology At Implantation Site
Histology from sheep #4 demonstrating the path of electrode (seen as round voids) as it tunneled into
the myocardium. There is moderate myocardial degeneration and necrosis with a mixed inflammatory
reaction including macrophages, lymphocytes and neutrophils.
164
Figure 7-7 Radiograph of Micropacemaker
Radiograph of a fetal sheep heart specimen fixed in formaldehyde (sheep #4). The electrode is
connected to the micropacemaker via an intact flexible lead.
X-Ray imaging of the micropacemaker system and surrounding tissues were
performed after necropsy in sheep #1, 4 and 5 in order to determine the integrity of
the flexible lead / electrode prior to complete removal of the hardware from the
tissue. A broken flexible lead was seen with the first device (fracture at the insertion
of the lead into the micropacemaker) and was determined to be secondary to a
weakening of the wire at that location during the fabrication process. This joint was
165
augmented for future experiments and there were no lead fractures for the
subsequent six experiments (Figure 7-7).
Discussion
CHB in the human fetus with hemodynamic compromise presents a significant
therapeutic challenge. While many fetuses with CHB can survive to birth, more than
a quarter of these fetuses will develop hydrops fetalis, resulting in nearly universal
mortality unless early delivery can be achieved.[40] Cardiac pacing is the standard
therapy for heart block, but a method to successfully pace a fetus has been elusive, as
no human fetuses that have undergone fetal pacing have survived. Our
micropacemaker system was designed to overcome one of the most important
hurdles of fetal pacing: the ability to place the entire pacing system completely inside
the fetus. Most previous fetal pacing attempts have placed electrodes on the fetal
myocardium with pacing wires exiting the fetal body into a pacemaker implanted in
the maternal abdomen.[34-37] The presence of this extracorporeal pacing wire is
problematic in that fetal movement may result in lead dislodgment or impingement
of other structures (including the umbilical cord). Recently, placement of a
conventional pacemaker into a fetus was attempted from a left fetal thoracotomy via
maternal laparotomy and hysterotomy (open fetal surgery).[56] Fetal demise was
seen on POD #5, and while the risk of surgery to the fetus is evident, the maternal
risks of open fetal surgery must also be considered.[62, 63] The high risks associated
166
with open fetal surgery were a primary impetus for developing a percutaneous
delivery system and a micropacemaker that was small enough to be implanted
completely within the fetus via instruments and approaches in routine use for other
minimally invasive fetal interventions.[64] Thus, our fetal micropacemaker
addresses the two major challenges of fetal pacing by 1) being a fully implantable
pacing system, and 2) allowing for percutaneous implantation without the need for
open fetal surgery.
Our initial human targets for fetal micropacemaker implantation are expected to
be those fetuses with CHB and hydrops fetalis. In the human fetus, we expect to target
the ventricles directly through the chest, as opposed to the subxyphoid approach used
in our fetal sheep model. These fetuses are expected to have pericardial and pleural
effusions, which should facilitate access to the fetal heart and implantation of the
micropacemaker. A major challenge of our experimental model was accurately
placing the cannula tip within the pericardium (past the parietal pericardium) and
directly against the epicardial surface. Advancing the cannula through the
pericardium without injuring the underlying epicardium (separated at times by only
a potential space) proved difficult and resulted in cardiac perforation in one case.
With realization that deployment of the device in a non-hydropic fetus was a
significant limitation of the fetal sheep model, we altered our experimental protocol
to include creation of a pericardial effusion by needle injection of saline into the
pericardial space. Although this pericardial effusion resulted in improved
167
visualization, appropriate placement of the needle tip in the pericardial space
remained challenging. In addition, concern for rapid creation of a pericardial effusion
resulting in acute cardiac tamponade tempered the amount of injected saline and
therefore limited the size of the effusion. While there are other models of fetal
hydrops,[65, 66] they are limited by significant hemodynamic compromise and fetal
morbidity. In addition, although fetal sheep models of CHB have been proposed,[67,
68] creating iatrogenic fetal heart block was not necessary, as we were able to clearly
identify ventricular capture (as well as ascertain the pacing rate) despite a faster
intrinsic fetal heart rate.
Because our pacing system does not allow for alterations in pacing outputs after
implantation, we chose an output based on the limited literature on pacing human
fetuses, as well as data from our rabbit experiments.[60] A single case report provides
a pulse-duration curve for a human fetal pacing[36] and is consistent with pacing
parameters seen with epicardial pacing in neonates, children and adults.[69] Results
of that experiment, as well as data from our acute electrode testing in the rabbit model,
however, did not predict the elevation of pacing thresholds over time in the
developing fetal lamb. As the literature suggests that fetuses exhibit less fibrosis than
children and adults,[70] we did not anticipate the observed degree of myocardial
inflammation and subsequent capture threshold elevations after electrode placement.
However, after observing the inflammatory reaction that was present histologically
at the electrode-myocardial interface, it is clear that a significant inflammatory
168
response can occur in the fetal myocardium. It is uncertain, however, what degree of
inflammation is due to mechanical stresses applied to the myocardium by electrodes
that were also attached to the pericardium rather than moving freely with the
epicardial surface as would be expected with proper electrode placement in a fetus
with a larger pericardial effusion. Due to these findings, we anticipate including a
steroid-eluting plug in our screw electrodes in the fetal micropacemaker device, as is
standard in commercially available pacemaker leads.
Our micropacemaker has a less than 3.5 mm diameter and is 5 to 7 times smaller
than the newly developed leadless pacemakers currently under investigation for
transvenous implantation [MicraTM by Medtronic, Inc. (Minneapolis, MN) and
NanostimTM by St. Jude Medical, Inc. (St. Paul, MN)]. The very small size (0.15 cc
volume) and cylindrical shape of our pacemaker necessitates a small battery with
very limited charge capacity. As a result, we designed and incorporated a recharging
system to allow the pacemaker device to function for the duration of the pregnancy
(anticipated 3-5 months). While the technical details of this recharging system are
reported separately,[61] it is worth noting that the electronic circuit was redesigned
during these experiments such that small fluctuations in the recharging field seen by
the micropacemaker would result in relatively large changes in the pacing rate. This
modification proved very useful for optimizing the recharging time since changes in
fetal and maternal positioning could be quickly recognized (by a decrease in the pacer
rate) and the recharging coil re-positioned accordingly.
169
During the follow-up period, three fetuses expired and were prematurely
delivered. Although the necropsy suggested that infection was the cause of death for
sheep #4, a cause of death was not determined for sheep #6 and #7. Good fetal health
during the follow-up periods had been suggested by ultrasound, suggesting that the
animals had tolerated the initial implantation procedures. The deaths were therefore
unexpected in both cases. An R-on-T phenomenon as the cause of a fatal ventricular
arrhythmia remains an unlikely possibility in the setting of VOO (fixed pacing) rates
that were slower than the intrinsic fetal heart rate in our model. This is not expected
to be as relevant to our intended human recipients since the device pacing rates of
>100 bpm will far exceed the intrinsic rates of the fetus with complete heart block.
The micropacemaker implantation procedures were significantly more invasive in
the fetal sheep model compared to the planned percutaneous insertion in a human
fetus. This was due to the need to surgically open the uterus for placement of fetal
ECG leads for confirmation of ventricular capture in the setting of rapid intrinsic fetal
heart rates. These more invasive procedures performed in the fetal sheep model may
have significantly contributed to the fetal demise.
The results of our acute and chronic fetal sheep testing are encouraging for the
viability of this system for human fetal use. Although the lack of a significant
pericardial effusion in the fetal sheep model resulted in a more complex implantation
procedure than we expect in a human fetus with hydrops fetalis, we were able to
successfully implant the electrode into the fetal myocardium and confirm ventricular
170
capture. In addition, we were able to recharge the device wirelessly through the
maternal abdomen. The main design risk that we identified before this study was
potential mechanical failure of the flexible lead between the epicardial electrode and
the micropacemaker device in the diaphragm / abdomen. After correction of a wire
joint design flaw recognized in the first experiment, we had no further evidence of
damage to the flexible lead in these implanted micropacemaker systems. Our device
has been granted a Humanitarian Use Device designation by the Food and Drug
Association, and we ultimately plan to progress to a clinical trial for device approval
under a Humanitarian Device Exemption.
Conclusions
A percutaneously-implantable fetal micropacemaker system provides a possible
new approach to pacing fetuses with CHB and hydrops fetalis. The device is
completely implantable within the fetus and thus minimizes the risks of lead
dislodgement due to fetal movement. In addition, the minimally invasive
percutaneous implantation approach avoids the risks of open fetal surgery. While the
fetal sheep model presented some unique challenges for implantation compared to a
hydropic human fetus, our data provide support for and demonstrate the necessity of
a clinical trial to confirm that the fetal micropacemaker is a viable option for
treatment of this potentially fatal condition.
171
Acknowledgements
We thank Dr. Joshua Kramer and Charter Preclinical Services for histologic
interpretations and analyses of the necropsy specimens. This research was funded
by an NIH R01grant (1R01HD075135), the Southern California Clinical and
Translational Science Institute, and the Coulter Foundation.
Clinical Perspectives
Successfully pacing a human fetus continues to be the “holy grail” of pacing.
Although a very small number of human fetuses require pacing, those needing it often
depend on pacing for survival due to the extremely high mortality rates associated
with profound bradycardia. Previous fetal pacing attempts have relied on
technologies developed for larger patients, and no survivors of chronic fetal pacing
have been reported to date. The fetal micropacemaker may represent a major
breakthrough in fetal pacing due to the development of a pacing device specifically
intended for human fetuses. The fetal sheep data described in this manuscript
demonstrates successful ventricular capture with the micropacemaker device but
also highlights the difficulties of the non-hydropic animal model. We anticipate
further animal implants with creation of a larger pericardial effusion in an adult large
animal model. Such data may provide sufficient justification to implant the first
human device under a compassionate-use scenario in a hydropic fetus with heart
172
block. Demonstration of successful pacing in a hydropic human fetus will ultimately
be required to demonstrate the efficacy and safety of the device.
173
Chapter 8: Assessing Cardiac Capture and Safety Factor
Adriana N. Vest, Li Zhou, Yaniv Bar-Cohen, Gerald E. Loeb
Preface
This chapter describes a method for determining safety factor for a given stimulus
strength and presents results from four of the seven fetal sheep from the study
presented in Chapter 7.
Abstract
We have developed a rechargeable fetal micropacemaker in order to treat
severe fetal bradycardia with comorbid hydrops fetalis, a life-threatening condition
in pre-term non-viable fetuses for which there are no effective treatment options. The
small size and minimally invasive form factor of our design limit the volume available
for circuitry and a power source. The device employs a fixed-rate and fixed-amplitude
relaxation oscillator and a tiny, rechargeable lithium ion power cell. For both research
and clinical applications, it is valuable to monitor the electrode-myocardium interface
in order to determine that adequate pacemaker output is being provided. This is
typically accomplished by observing the minimal stimulus strength that achieves
threshold for pacing capture. The output of our simple micropacemaker cannot be
programmatically altered to determine this minimal capture threshold, but a safety
174
factor can be inferred by determining the refractory period for ventricular capture at
a given stimulus strength. This is done by measuring the minimal time interval
between naturally occurring QRS complexes and successful stimuli. The method was
tested in a pilot study in four fetal sheep and the data demonstrate that a relative
measure of threshold is obtainable. This method provides valuable real-time
information about the electrode-tissue interface.
Introduction
Progressive complete heart block with comorbid hydrops in the human fetus is a
rare but life-threatening condition.[28, 30, 38] Once hydrops develops as a result of
heart block, fetal demise is nearly inevitable if the fetus cannot be delivered due to
prematurity or other clinical concerns; several hundred fetal deaths a year are
attributable to this condition.[29] Successfully pacing the fetus would be expected to
restore adequate blood flow and allow resolution of hydrops fetalis within several
weeks, permitting an otherwise normal gestation and delivery. After birth, the infant
could then be implanted with a standard pacemaker with epicardial leads. We have
developed a fetal micropacemaker to meet this prenatal need and are currently
performing pre-clinical studies.
We can better understand the requirements for a fetal pacemaker by considering
the failure modes from previous attempts to pace fetuses, all of which have been
175
unsuccessful and ended with fetal death.[35-37, 55, 56] The causes of death were not
clearly identified in all cases, but were likely due to three possible etiologies:
Surgical complications from open surgery [55, 56]
Lead placement complications [35, 36]
Lead dislodgement [35] or strangulation
These failure modes suggest two major design requirements for successful fetal
pacing:
A minimally invasive technique to reduce fetal and maternal
morbidity due to surgical complications from an open or elaborate
procedure
A device that can be wholly contained within the fetal chest,
avoiding the need for a trans-uterine lead.
The device that we have developed addresses both of these requirements by
utilizing a percutaneous approach through a fetal surgical cannula and a packaging
scheme that can fit entirely in the chest wall. To realize a percutaneous approach
using a standard fetal surgical cannula, the device must be in the form of a cylinder
that will fit through the cannula’s inside diameter of 3.8 mm. Because the market size
is less than 500 devices a year, the technology must be simple and inexpensive both
to develop and to build.
176
To engineer such a device, we employ a tiny rechargeable cylindrical battery and
minimal circuitry. Battery capacity is proportional to volume, and in this case, reaches
a maximum of 3mAh at full charge. In order to meet stringent requirements for size,
low power consumption and low development cost, we have utilized a simple
relaxation oscillator based on a single transistor.[60] The rate and output strength of
the oscillator is set by component values during fabrication but varies also with
supply voltage. This relationship is calibrated during manufacture of each
micropacemaker; cell voltage and output pulse strength is readily determined from
the pacing rate at any time. The 3 mAh cell can sustain a typical stimulus output pulse
of 3 µC (3V peak, 250µs time constant) for 6 days and can then be recharged by
inductive coupling of a 6.78MHz electromagnetic field from a transmitting coil
positioned outside the maternal body.
The simple electronic circuit leads to several consequences in the design. There is
no data communication with the implant and its nominal pacing rate and output
strength must be pre-set during fabrication. It is desirable to set the stimulus strength
as low as possible to conserve power in order to maximize the recharge interval, but
it is also important to include a safety factor to ensure effective ventricular capture
for somewhat unpredictable electrode placements and tissue conditions.
A notable drawback to the minimal circuitry is the inability to change the stimulus
strength if capture thresholds rise above the programmed pacing output. In addition,
177
in order to measure actual capture thresholds, varying the strength of the stimulus
(amplitude, duration or both) is typically required to determine the minimal strength
necessary to elicit a contraction (capture). Since capture thresholds can change over
time, an ongoing ability to determine capture thresholds is especially important for
monitoring the electrode-myocardium interface for the following:
healing and formation of scar tissue around the electrode;
mechanical stress on the tissue; and
lead dislodgement.
While these changes can be seen on necropsy, it is important to determine when
such events occur in real time in order to better predict pacing or understand failure
modes.
Although a typical strength-duration capture curve is not attainable with the fetal
pacemaker, it is possible to infer a relative measure of threshold in a constant output
scenario by measuring the minimal interval necessary between an intrinsic
contraction of myocardium and a stimulus in order to achieve successfully paced
ventricular contraction. While the interval between naturally occurring contractions
and pacing stimuli could be varied systematically by detecting an intrinsic heart beat
(QRS signature) and then commanding the stimulation to occur at a certain delay
after that beat, this is not possible with our simple micropacemaker. Instead we used
a fixed pacing rate that was lower than the intrinsic fetal heart rate, which results in
random time intervals between intrinsic beats and subsequent pacing stimuli,
178
ensuring that a large number of possible intervals occur eventually. By plotting
successful (X) versus non-successful (O) intervals, a boundary emerges that is here
defined as the minimal interval (MI) and is an estimate of pacing refractory periods
(and hence safety factor) (Figure 8-1, A). The minimal interval and the output
strength used to generate it correspond to a point on the strength interval (SI) curve
(Figure 8-1, B).
Changes in the ability to elicit a contraction of the myocardial tissue would
become evident in the strength-interval curve. It is well understood that pacing
threshold tends to rise over time due to the normal inflammatory process of the
foreign body reaction, rapidly in the acute phase and then gradually leveling out to a
value somewhat higher than implant day.[71, 72] This apparent rise in threshold is a
matter of increasing distance and lower charge density experienced by the nearest
excitable myocytes. More strength is necessary to overcome these factors and the SI
curve would appear to shift up (Figure 8-1, C). A similar change would result from
any increase in the distance between the electrode and the excitable myocytes, such
as injury and cell death around the electrode tip or lead dislodgement. The shift up
would be evident by a lengthening in the minimal interval measured at a given
strength. A change in threshold of the individual myocytes would also shift the SI
curve. If autonomic sympathetic tone increased, as it does during a stress response,
the contractility of the myocytes would increase and the apparent threshold of the
tissue would decrease, causing a shift in the strength interval curve downwards. This
179
would allow a given stimulus strength to capture at a shorter minimal interval.
Because changes in autonomic tone generally result also in a change in intrinsic heart
rate, it should be possible to distinguish these two mechanisms.
Figure 8-1 Theory of Minimal Interval
The boundary observed in (A) and its associated interval and strength make up a point on the strength
interval curve (B). Changes in the ability to elicit a contraction of the myocardial tissue would become
evident in the strength interval curve as an upwards or downwards shift (C) and would appear as a
longer or shorter minimal interval, respectively. If diastolic threshold is estimated, the safety factor of
a given stimulus strength could be determined (D).
Minimal interval measurements by themselves provide a method to infer the
relative threshold of the tissue over time, but they could also be used to generate a
more complete picture of the threshold of the tissue and the safety factor if at least
180
some variation of stimulus strength could be achieved. Our micropacemaker does not
allow programmatic control of stimulus strength, but as the supply voltage from the
lithium cell drops there is a corresponding drop in output charge, which can be
measured accurately due to calibration during the manufacturing process. Allowing
the supply voltage to drop, noting output charge, and making minimal interval
observations can lead to the determination of many points on the strength interval
curve, assuming that the electrode interface and the autonomic tone are steady
during this period. The supply voltage can be increased quite rapidly as well by
recharging the device to the desired level; full recharge takes less than three hours
with a favorable placement of the external transmitting coil.[61] The diastolic
threshold (DT) can be estimated from such a calibrated strength-interval curve and
the absolute safety factor of stimulation can be determined (Figure 8-1, D).
This safety factor measurement would be limited by the range of supply voltages
achievable between the maximal safe voltage for the lithium cell (4.0 VDC) and the
minimal supply voltage at which the relaxation oscillator generates output (1.8 VDC).
Additional supply voltage is generated during the recharging process, and the newest
version of the pacemaker can achieve an increase in the voltage supply of 1V per
applied mA in recharging current (up to 3 mA maximal safe charging rate).[73] This
enlarges the range of possible stimulus strengths, but the relationship is noisy as a
result of the effects of the high field strength from the 6.78 MHz recharging coil (note
inconsistent minimal intervals in Figure 8-6 during recharging).
181
Method
The animal study protocol was approved by the Institutional Animal Care and Use
Committees at the University of Southern California and the Los Angeles Biomedical
Research Institute. Fetal micropacemakers with rates between 100 and 150 bpm
were implanted into seven fetal sheep according to the procedure discussed in Bar-
Cohen et al.[57] In 6 of 7 experiments, the micropacemakers had a 3 µC output pulse
on the day of implant (3 V peak, 250µs time constant). With the 6th implant, a higher
output device (6.1 µC; 2-day recharging interval) was used. Devices were implanted
fully charged, but supply voltage and therefore output charge decreased over time.
Devices were recharged in some cases, which increased supply voltage and output
charge. Implants from sheep 4 through 7 produced adequate data for the analysis
presented here.
The fetal electrocardiogram was detected by three trans-uterine electrodes
attached directly to the fetus. The electrodes were sewn directly onto the chest
surrounding the heart instead of on the extremities to avoid shunting the signals by
the amniotic fluid. The fetal skin leads extended out of the maternal abdominal
incision could be placed into a pouch sewn on the maternal skin surface. These fetal
skin electrodes were connected to an external biosignal monitoring device (BioRadio
150, Great Lakes NeuroTechnologies) and the digitized signal was captured on PC,
filtered to eliminate DC drift and high frequencies (LPF=400 Hz, HPF=5 Hz). The
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signals were subsequently analyzed, and each interval between the preceding
intrinsic QRS signature and the subsequent stimulus artifact was measured and
assessed for capture. Successful capture was detected by the presence of a
prematurely occurring ventricular contraction immediately after the pacing stimulus.
Successful (X) vs. non-successful (O) stimuli were then plotted as a function of the
interval between the previous QRS complex and the stimulus artefact (Figure 8-1, A).
The duration of ECG analyzed depended on the availability of data and the length of
data necessary to generate a clear minimal interval boundary, generally from 15 to
150 seconds. The boundary between the symbols denotes the minimal interval for
successful capture and corresponds to the ventricular capture refractory period.
Each implantation built on the lessons learned from previous experiments, and
therefore each was unique. The resulting observations and interpretations will be
presented for each case along with a summary of observations provided in Table 8-1.
Results
Sheep 4
Observations
Measurements made on operational day (OD) indicate a minimal interval of
0.155s. A clear boundary between successful and unsuccessful stimuli was not
183
observed, and the minimal interval was taken to be the earliest successful stimulus.
This assumption was made for all subsequent analyses. Measurements made on post
op day 1 (POD1) indicate that the minimal interval lengthened somewhat (0.171s vs.
0.155s) but the overlap region was reduced. On POD3, measurements were again
taken yielding a minimal interval of 0.109s, much lower than OD (0.155s) or POD1
(0.171s). Upon necropsy, the pacemaker was observed to have been implanted in a
nearly perfect location, with the battery case wedged into the diaphragm, the flexible
lead partially extended and straight, and the electrode buried in the left ventricular
wall.
Figure 8-2 Minimal Interval for Sheep 4
Thresholds for Sheep 4 were seen to lengthen on POD1 from OD, and then drastically shorten on POD3,
indicative of fetal distress.
184
Interpretation
Because the pacemaker was implanted in a very favorable location, the cardiac
tissue threshold at implant day and therefore the measured minimal interval of
0.155s was most likely an indication of a very low pacing threshold. The variability in
threshold indicated by the substantial overlap region on OD could have been related
to local irritability from bleeding or other tissue damage around the electrode.
Because the stimulus strength was nearly constant between OD and POD1 (2.77 µC
vs 2.70 µC), and no obvious autonomic effects were evident in the stable heart rate,
the threshold of the tissue likely increased slightly, a response consistent with the
beginning stages of electrode stabilization and the inflammatory response.
On POD3, the much shorter minimal interval measured was unexpected. It was
likely due to autonomic affects resulting from stress to the fetus as indicated by a
substantial increase in the intrinsic fetal heart rate (262 bpm vs 178 bpm). This is
consistent with the spontaneous abortion of the fetus just 2 days later on POD5 (Table
8-1).
Sheep 5
Observations
A pacemaker was implanted in Sheep 5 (device I, FP1D02C041003A) and when
this device did not show capture immediately, a second device was implanted (device
185
II, FP1D02C041013A). After several minutes, both devices started capturing
occasional ventricular contractions. The minimal intervals on OD measured for
devices I and II were 0.151s and 0.121s, respectively. On POD2, device I was no longer
capturing the tissue, but device II was still capturing the tissue and with a longer
minimal interval period of 0.161s (vs. 0.121s) and less overlap observed. This
measurement was in the presence of a lower heart rate (about 188 bpm instead of
242 bpm).
At necropsy, device II was observed to be well attached to the heart with several
turns of the electrode corkscrew embedded in the epicardium, while device I was
instead adjacent to the myocardium. The second pacemaker electrode implanted,
device II, was actually screwed through the first electrode, putting pressure on it in
the direction of the epicardium (Figure 8-4).
186
Figure 8-3 Minimal Interval for Sheep 5
Implantation of two devices into Sheep 5 complicated results, but it was shown that device II had a
better electrode-tissue interface.
Figure 8-4 MicroCT on Sheep 5
This microCT image taken of the cardiac tissue blocked around the implanted electrodes shows the
orientation of the electrodes as they were implanted into the epicardium. Electrode II is interwound
in electrode I.
187
Interpretation
The lower minimal interval seen for the second device on OD suggests that this
device had a more favorable electrode tissue interface. Considering that the minimal
interval measurements from both devices were subject to the same autonomic effects
and both devices generated the same stimulus charge, the difference between the
measurements must be entirely due to the electrode tissue interface. This
interpretation is further confirmed by the necropsy results where it was observed
that device II was screwed through device I and well anchored into the myocardium,
while device I was not (Figure 8-4).
Capture was seen on POD2 for device II only and with a longer minimal interval
than what was observed on OD (0.161s vs 0.121s). This change in interval could be
due to a reduction in autonomic sympathetic tone and an increase in threshold of the
myocytes (indicated by a drop in heart rate), or from mechanical effects of the
intertwined electrodes, which would have applied unknown stresses to the
myocardium, damaging the tissue and perhaps lengthening the distance between the
active electrode tip and healthy excitable myocytes. The stimulus strength also
dropped from OD to POD2 as a result of gradual discharge of the lithium cell, which
could also be a contributing factor to the lengthening of the minimal interval.
188
Sheep 6
Observations
For Sheep 6, a pacemaker with a larger stimulus output (6.1 µC vs 2.7 µC) was
used. This was chosen because it was not yet clear that the inability to capture for
longer than 2 days in Sheep 5 was due to poor electrode placement (necropsy and
histology had not yet been completed). The higher stimulus output resulted in a faster
battery depletion, so recharging took place on POD2. Capture of the heart tissue was
seen on OD after implant and on POD2 after recharging the device. On OD, the minimal
interval observed was 0.175s. Capture was not seen between 0.2 and 0.3s intervals
because the capture interval (0.384s; rate period = 0.405s) and intrinsic rate 157 bpm
(period = 0.382s) were so similar, resulting in long segments of continuous capture.
The device continued to capture the myocardium on POD2, showing a slightly shorter
minimal interval of 0.168s. Heart rate on POD2 was fluctuating between 170 and 200
bpm during measurement, as opposed to the stable heart rate of about 157 bpm seen
on OD. Upon necropsy, the electrode was touching the myocardial surface, but was
not engaged into the myocardium.
189
Figure 8-5 Minimal Interval for Sheep 6
Intervals for Sheep 6 show a similar duration to those of Sheep 4. However, the pacemaker implanted
in Sheep 6 had a much higher stimulus output, and therefore threshold is estimated to be much higher
than it was in Sheep 4.
Interpretation
Considering that the output stimulus charge used for Sheep 6 was over double the
intensity of the other three sheep, thresholds inferred in this experiment are not able
to be compared directly with Sheep 4, 5, and 7. However, we can make a general
conclusion that because the minimal intervals measured for Sheep 6 were about the
same as previous implants while the output strength used was much higher,
thresholds were probably much higher than Sheep 4 and 5. This would be consistent
with the poor electrode tissue interface confirmed upon necropsy and histology.
Nevertheless, the pacemaker continued to capture the myocardium on POD2, most
likely because of the very high output stimulus used. The slightly shorter minimal
190
interval was probably related to transient autonomic elevation of the fetal heart rate
on POD2.
Figure 8-6 Minimal Interval for Sheep 7
Thresholds of Sheep 7 were higher than Sheep 4 and 5, indicating an inferior electrode-tissue interface,
which was confirmed with necropsy and histology.
191
Sheep 7
Observations
On OD for Sheep 7, a low output stimulus device was implanted and the fetal heart
rate was very stable around 177 bpm. Capture was only seen on OD and was not able
to be restored on POD2. Recharge of the pacemaker was performed to attempt to
raise the output and capture the tissue, but was not successful. Therefore,
measurements are only included for OD.
The minimal intervals early in the implant were around 0.200s and settled to
about 0.194s at the end of data collection, indicating little change over time. The
recharging equipment was tested during this implantation procedure and minimal
interval was measured during recharging to determine if the increased stimulus
strength from the higher supply voltage had an effect. A minimal interval of 0.189s
was observed, but capture was inconsistent at intervals longer than 0.189s (Figure
8-6, OD Recharging). At necropsy of the implant, the electrode was easily detached
from the myocardium and the friction disk was seen to lie on the outside of the
pericardium
192
Interpretation
The evidence from necropsy indicates poor contact between the electrode and
myocardium, which is supported by a very long minimal interval in comparison to
good electrode insertions seen in earlier animals (sheep 4 and 5).
The decreased minimum intervals during recharging (0.189s vs 0.200s, Figure 8-6,
OD Recharging) could be attributed to the higher charge delivered to the tissue during
recharging. The recharging process induces a higher potential at Vsupply,
accompanied by a ripple due to the imperfect nature of the half wave rectification
circuit used to recharge. This higher Vsupply also raises the gate voltage of the PUT
thereby increasing the stimulus amplitude. Despite a decrease in minimal interval,
many unsuccessful stimuli were seen above the minimal interval and substantial
overlap was observed. The reason for this overlap is unclear, but may be a result of
the instability of the electrode-tissue interface from insertion trauma. Another
potential cause could be the variation in amplitude of pulses, a consequence of the
noise in the relationship between the recharging current and output charge.
On POD2, no device capture was seen, even during recharging. This indicates that
the thresholds had risen and the poor electrode tissue interface could not be
overcome.
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Table 8-1 Summary of Results
This table presents the devices implanted into four fetal sheep, along with their follow up details. It
includes the stimulus charge delivered, the minimal interval, and the intrinsic fetal heart rate.
Discussion
The pilot study presented here of indirectly assessing the threshold of cardiac
tissue showed that minimal interval observations convey valuable information about
the status of each implant. The variety of electrode placements and clinical outcomes
in these case studies provided a variety of data for interpretation. The results of our
minimal interval analysis were generally consistent with the conditions observed
post mortem. A poor or deteriorating electrode tissue interface can be distinguished
Sheep Day Device
Battery
Voltage
Charge
Dataset
Length
Minimal
Interval
FHR
(bpm)
Summary Termination
Op Day 3.9V 2.77µC 146sec 0.155sec 165-210
POD1 3.8V 2.70µC 60sec 0.171sec 178
POD3 3.6V 2.56µC 60sec 0.109sec 262
Device I
FP1-D02C041003A
3.9V 2.73µC 36sec 0.151sec 242
Device II
FP1-D02C041013A
3.9V 2.73µC 36sec 0.121sec 242
POD2
Device II
FP1-D02C041013A
3.7V 2.60µC 15sec 0.161sec 180-196
OD 3.9V 6.10µC 110sec 0.175sec 157
POD2 3V 4.48µC 30sec 0.168sec 170-200
Early 3.7V 2.67µC 36sec 0.200sec 177
Late 3.7V 2.67µC 28sec 0.215sec 174
Recharge 3.7V 2.67µC 50sec 0.189sec 174
Post-recharge 3.7V 2.67µC 16sec 0.194sec 178
4
5
6
7
FP1-ED173025A
Devices implanted tangential to the myocardium and with
electrodes paired. 003 implanted first and thought to have
been pushed into myocardium with 013. Xray confirms. In
sheep 5, there was a clearly visible electrode path in the
area of electrode 2 that extended approximately 3.7mm
from the epicardial surface.
The pacemaker was perfectly located, with the nose
wedged into the diaphragm, the flexible lead partially
extended and straight and the electrode buried in the left
ventricular wall.
Electrode tip may have been up against the epicardium but
was clearly not engaged in myocardium (confirmed by
path report); deformation seen from pressure of electrode
tip pushed against epicardial surface by tension in
pericardium (most likely). The location of the electrode
outside the myocardium is consistent with the high and
increasing threshold, similar to our previous implant, so
myocardial fibrosis was probably not the cause this time.
The electrode may have been slightly into the epicardium,
but the disc was on the outside of the pericardium which
means that the cannula probably did not penetrate through
the pericardium. POD 2 No Capture, even after recharging.
POD5 - Sheep
Aborted fetus
FP1-D06G047002A
FP1-D06G047007A
Op Day
Elective
termination
POD15
POD5 - Pre-
mature labor
and miscarriage
Miscarriage POD
5
194
from an excellent one. For instance, the device implanted in sheep 7 had a much
higher minimal interval than the device implanted in sheep 4 (MI = 0.155s vs 0.200s).
Necropsy showed that the electrode in sheep 4 was well embedded in the tissue and
the electrode in sheep 7 was not.
Data obtained from the four animal implantations were not adequate to establish
measures of safety factor, which would have required many follow up measurements
after the implant and fetus stabilized. Premature animal deaths, likely due to the
implantation procedure of fetal electrodes, meant relatively short follow-up on each
animal, and therefore limited time was available for exhaustive minimal interval
measurements.
The method employed in this paper to assess the capture of cardiac tissue has
several limitations, including the ability to control confounding variables, the time it
takes to collect data, and the limited situations in which this is a viable measurement
method. In order to determine if threshold changes are due to the electrode-tissue
interface, the effects of cardiac autonomic tone must be controlled or compensated.
This can be done by using heart rate as a measure of autonomic tone and only
comparing data obtained during comparable heart rates. Drugs that affect
sympathetic tone and fetal heart rate, such as positive or negative chronotropes,
could also be administered via the maternal circulation in order to provide direct
control of this variable during measurement. Data collection itself is a time-
195
consuming process, and the length of ECG needed to produce interpretable data can
range from a fraction of a minute to several minutes. This further compounds the
problem of controlling for autonomic effects because they can vary from second to
second, especially if the mother is not sedated. The stimulus strength in these
experiments was not held constant from day to day, and therefore added a variable
to consider when interpreting data. Stimulus strength could be controlled by starting
each data acquisition session with a recharging step, so that the power supply, and
therefore the charge delivered, from day to day is constant. The initial design of the
study included this step, but technical difficulties with the recharging system limited
its implementation in the studies presented here. Once this step is taken, interpreting
shifts in minimal-interval in terms of the desired strength-interval curve will be more
straightforward.
This method is only available as a measurement tool when the natural heart rate
is above the rate of the pacemaker, therefore restricting its use to pre-clinical animal
studies in healthy fetuses that have a high resting heart rate. A very good timing signal
of the fetal ECG is also necessary, which we obtain here using chest leads directly on
the fetus, a step that would not be feasible in human clinical studies. It has been shown
that obtaining fetal ECG from the maternal abdomen is possible, but the fetal signals
can at times be so attenuated or contaminated with other higher amplitude maternal
signals (i.e. maternal ECG or EMG) that the fetal QRS signatures are unreadable.[74,
75] The fetal ECG from the maternal abdomen would be adequate to confirm
196
successful fetal ventricular contractions (QRS signatures) and measure minimal
intervals as long as the placement of abdominal electrodes and fetal presentation
provide a good fetal signal-to-noise ratio. The stimulus artifact, which is much larger
in amplitude than the fetal ECG, should be readily detectable on the maternal
abdomen. Capture or non-capture is readily determined by Doppler ultrasound, but
the precision of timing of the preceding beat will be much less than from the QRS
complex of the fetal ECG.[76]
Observations on more implantations would need to be performed in order to
more fully characterize this novel method of assessing threshold and safety factor.
Additional data could include tracking a longer term device implant and then looking
for the typical threshold changes expected with pacemaker leads: a low threshold
during implantation, including short-term instability due to initial tissue injury, and
then a stabilization of threshold at a somewhat higher value over the ensuing
month.[71, 72] Future studies could further characterize the relationship between
thresholds and minimal interval, and prove the effectiveness of the measurement by
improving the quality and length of data collection.
In conclusion, this indirect method of measuring threshold is able to show
changes that result from autonomic effects and electrode-tissue interface. It yields a
quantitative estimate of threshold when stimulus strength cannot be varied
programmatically.
197
Acknowledgments
The authors would like to thank Drs. Ramen Chmait, Michael Silka, and Jay Pruetz
for implanting devices and their clinical expertise, and consultant Ray Peck for
fabrication and engineering expertise. This work was generously funded by the
Coulter Foundation, the NIH (grant #: 5R01HD075135-02), the Wright Foundation,
and Southern California Clinical and Translational Science Initiative.
198
Chapter 9: Conclusions and Future Directions
Adriana N. Vest
The results of this research show that the fetal micropacemaker has been
developed to meet all specified electrical requirements of the clinical indication,
including the provision of an adequate stimulus, a method of measuring that safety
factor in vivo during pre-clinical studies, and an ability to recharge the device
repeatedly over the device lifetime. The micropacemaker provides a pacing device
that can be implanted with a minimally invasive approach, a feature that could be
useful to a wider patient population that cannot undergo surgery to have a pacemaker
implanted. Therefore, future work will involve expanding the technology to be used
in adults and small children who have a need for a pacemaker that can be implanted
in this manner.
Adults and small children who can tolerate a minimally invasive procedure, but
not open-chest surgery, should still be able to tolerate a larger delivery cannula than
the one used in fetal surgery. Therefore, the new version of the micropacemaker can
be substantially larger and incorporate several features that will ensure higher levels
of functionality and longevity:
• a hermetic package;
• a larger lithium ion cell; and
199
• a custom chip with integrated communication.
It is important that a pacemaker for adult or pediatric patients have these features
because it will be expected to endure for years in the body. In addition to the novel
implantation procedure and small form factor, another advantage of this device
would be its ability to be recharged, a feature that is not currently offered in the
pacemaker or ICD market today. Such a rechargeable device could outlive its
surgically implanted counterparts, but only if the lead between the myocardial
electrode and the hermetic package can be engineered to survive the large number of
small movements imposed by each cardiac cycle (~40 million/year).
200
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Abstract (if available)
Abstract
A fetal pacemaker can dramatically improve the outcome for fetuses that develop complete heart block in utero. We have developed a rechargeable fetal micropacemaker in order to treat severe fetal bradycardia with comorbid hydrops fetalis. Implanting our fetal micropacemaker could reverse the typically fatal outcome of this condition, resulting in the resolution of hydrops within one to two weeks by pacing the heart and restoring adequate blood flow to the fetus. ❧ The main requirements of the device are that it be implanted with a minimally invasive technique and be implanted entirely within the fetal chest. The pacemaker developed meets these requirements by being designed to fit within a standard fetal surgical cannula, putting a hard constraint on the size of the device. The size limitation dictates the volume available for circuitry and a power source inside the implant. A simple fixed-rate and fixed-amplitude relaxation oscillator based on a single transistor provides stimuli while also meeting stringent requirements for low power consumption and low development cost. A commercially available cylindrical, rechargeable 3 mAh lithium ion cell provides power. ❧ In this dissertation, the limits of the simple circuitry and power source are identified and compensated for. A power budget provides an analysis of the battery life with any given combination of components. The main draw of current from the battery is the output pulse, so it is desirable to set the stimulus strength as low as possible to conserve power in order to maximize the recharge interval, but it is also important to include a safety factor to ensure effective ventricular capture for somewhat unpredictable electrode placements and tissue conditions. The unknown conditions of each electrode placement leads to a need to monitor the electrode-myocardium interface in order to determine that adequate pacemaker output is being provided. This is typically accomplished by observing the minimal stimulus strength that achieves threshold for pacing capture. The output of the micropacemaker cannot be programmatically altered to determine this minimal capture threshold, but a safety factor can be inferred by determining the refractory period for ventricular capture at a given stimulus strength. This is done by measuring the minimal timing between naturally occurring QRS complexes and successful stimuli. Upon pilot testing this method in four fetal sheep, data demonstrate that a relative measure of threshold is obtainable, providing valuable real-time information about the electrode-tissue interface. ❧ Limited volume inside the implant also constrains the engineering of the recharging system, which relies on inductive coupling to provide wireless current to the lithium ion cell. To overcome the lack of regulation circuitry within the implant, a method for controlling the recharging process was developed and utilizes pacing rate as a measure of battery state, a feature of the relaxation oscillator used to generate stimuli. The verification of the recharging system shows successful generation of recharging current in a fetal lamb model.
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Vest, Adriana Nicholson
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Electronics design and in vivo evaluation of a wirelessly rechargeable fetal micropacemaker
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Biomedical Engineering
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