Electromagnetic Suspension System for Prosthetic Limbs that Compensates

for Residual Limb Shrinkage

A THESIS SUBMITTED TO THE GRADUATE DIVISION OF THE UNIVERSITY

OF HAWAIʻI AT MĀNOA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS

FOR THE DEGREE OF

MASTERS OF SCIENCE

IN

MECHANICAL ENGINEERING

MAY 2014

By

Diane Bautista

Thesis Committee:

Scott Miller, Chairperson

Timothy Roe

Reza Ghorbani

ii

We certify that we have read this thesis and that, in our opinion, it is satisfactory in scope

and quality as a thesis for the degree of Master of Science in Mechanical Engineering.

THESIS COMMITTEE

_____________________________________________

Chairperson

_____________________________________________

_____________________________________________

iii

© Copyright 2014

By

Diane Hanauhope Bautista

iv

Acknowledgements

I would like to thank Dr. Scott Miller, Dr. Timothy Roe, MD, and Dr. Kai

Newton for their guidance, support, and patience throughout this journey. I would also

like to thank Dr. Reza Ghorbani his guidance and patience, as well as Sladjan Lazarevic

for helping me with various aspects of the design and testing of this project. Thank you to

the University of Hawaiʻi at Mānoa and the Robotics Laboratory for allowing me to use

their facilities for this project. Finally, I would like to thank my family and friends for

their support. I especially would like to thank my father for helping me with the

construction part of this project. Without the help and support from everyone, I would

not be in the position I am today. Working with all of you has been a pleasure and a great

learning experience that I will never forget. Thank you very much.

v

Abstract

When an amputation of a lower limb occurs, forces due to walking are absorbed

by the soft tissues of the residual limb. Due to the soft tissues of the limb no being

accustomed to handling such high forces, one of the major issues that occurs is residual

limb shrinkage. As the amount of pressure fluctuates, so does the balance between the

fluids in the body. Shrinkage of the limb can lead to many issues such as damage to the

limb and an uncomfortable socket fit. Currently, vacuum suspension is the leading

method to compensate for limb volume changes. However, vacuum suspension has many

flaws as well. There is a need for a new prosthetic suspension system with the capability

to adjust and control the pressure at different points in the prosthetic.

In this study, a new magnetic suspension system for lower limb prosthetics was

designed and tested. This magnetic suspension system for lower-limb prosthetics uses

electromagnets to control the amount of force exerted on the residual limb. When the

electromagnets are turned on, they will attract the material on the liner, thus creating a

secure attachment. The design incorporates an Arduino microprocessor and a motor

driver to control the amount of negative pressure by the electromagnets. In this study, the

physical model of the socket was prototyped and the strength of the electromagnetic

suspension system was tested in an Instron testing machine.

An amplified circuit was designed to control the electromagnetic attractive force

based on the output of force sensors attached inside the socket. As the pressure decreased

in the socket due to limb shrinkage, the electromagnets were programmed to inversely

increase attractive force. The Arduino microprocessor and motor driver were used to

adjust the power delivered to the electromagnets. This feedback loop was tested to ensure

proper function.

After reviewing all the results, it was found that an array of electromagnetics

provided adequate force for attachment of the socket to the liner. The system was also

capable of varying the attractive forces at different points based on the force sensor

output. This research sets a completely new direction in the area of prosthetics, forever

changing what we already know about prosthetics.

vi

Table of Contents

Acknowledgments iv

Abstract v

Nomenclature viii

List of Tables ix

List of Figures x

1. Introduction 1

1.1. Amputation: An Overview 2

1.2. Biomechanics of Lower Extremity Sockets 6

1.2.1. Shear, Slippage, and Friction 7

1.2.1.1. Tissue Responses to Mechanical Loading: Normal 8

and Shear Force

1.2.1.2. Slippage 10

1.2.1.3. Frictional Forces 11

1.3. Interstitial Fluid 12

1.3.1. Limb Shrinkage 13

1.3.2. Physiology of Residual Limb Shrinkage 16

1.4. Prosthetic Devices 17

1.4.1. History of Prosthetics 17

1.4.2. Modern Prosthetic Limbs 19

1.4.2.1. Types of Prosthetic Suspension Systems 21

1.4.2.1.1. Sleeve Suspension 21

1.4.2.1.2. Pin and Lock Suspension 22

1.4.2.1.3. Suction Suspension 23

1.4.2.1.4. Vacuum Suspension 24

1.4.2.2. Types of Prosthetic Sockets 27

1.4.2.2.1. Patellar Tendon Bearing Socket 27

1.4.2.2.2. Total Surface Bearing Socket 28

1.4.3. Leading Prosthetic Companies 29

1.5. Literature Review 31

1.5.1. Residual Limb Volume 31

1.5.2. Comparison of Socket and Suspension Types 31

1.6. Research Motivation 33

1.6.1. Contributions by ME 696 Group 35

1.6.1.1. Liner Design 36

1.6.1.2. Socket Design 37

1.7. Thesis Organization 39

1.7.1. Design and Prototype of the New Electromagnetic 40

Suspension System

1.7.2. Development of the Feedback Control of Electromagnets 40

by Pressure Sensor Output

1.7.3. Mechanical and Functional Testing 40

2. Concept and Design of Electromagnetic Suspension 41

2.1. Introduction 41

vii

2.2. Socket Design 42

2.3. Liner Design 44

2.4. Power Supply 48

3. Mechanical Performance 50

3.1. Introduction 50

3.2. Instron Testing Setup 50

3.3. Results and Discussion 53

3.4. Conclusions 55

4. Functional Testing 56

4.1. Introduction 56

4.2. Control of Electromagnets 56

4.2.1. Tekscan Force Sensors 57

4.2.2. Arduino Microprocessor 60

4.2.3. Rover 5 Motor Driver 62

4.3. Experimental Setup for Testing the Circuit Operation 65

4.4. Results and Discussion 67

4.4.1. Arduino Output 67

4.4.2. Motor Driver Output 69

4.5. Conclusions 71

5. Conclusions and Future Work 72

5.1. Summary 72

5.2. Conclusions 73

5.3. Contributions 74

5.4. Future Work 74

References 77

viii

Nomenclature BK: Below-Knee

GRF: Ground Reaction Force

PTB: Patella Tendon Bearing

TSB: Total Surface Bearing

US: United States

ix

List of Tables

Table 1.1: Age-Specific Estimates of Prevalence by Sex, Race, and Ethnicity 3

Table 1.2: Estimates of Prevalence by Type and Level of Limb Loss and Etiology 4

Table 1.3: Projected Prevalence of Limb Loss by Etiology and Age 5

Table 4: Force Required to Pull Apart – Electromagnet/Permanent Magnet02 46

(Settings: 3.50 PC/MU, 10 MU/Volt)

Table 5: Force Required to Pull Apart – Electromagnet/Steel Bar02 47

(Settings: 3.50 PC/MU, 10 MU/Volt)

x

List of Figures

Fig. 1.1. Projected Number of Americans (1000s) 6

Living with Limb Amputation Secondary to Dysvascular

Disease (Based on a Hypothetical Change in Incidence on

Amputation post-2005)

Fig. 1.2. Diagram of Normal and Shear Stress Applied at a Certain Point 9

on the Socket

Fig. 1.3. Interstitial Fluid 13

Fig. 1.4. Pressure Distribution between Blood Vessel and Interstitial Fluid 13

Fig. 1.5. Prosthetic Toe Found in Egypt 19

Fig. 1.6. Typical Trans-Tibial Prosthetic 20

Fig. 1.7. Sleeve Suspension 22

Fig. 1.8. Pin and Lock Suspension 23

Fig. 1.9. Suction Suspension 24

Fig. 1.10. Vacuum Suspension 26

Fig. 1.11. Cross section of vacuum suspension system. 26

Fig. 1.12. Patellar Tendon Bearing Socket 28

Fig. 1.13. Comparison of Electromagnetic Suspension System 35

and the Current Problem

Fig. 1.14. Liner design from ME696 student team 37

Fig. 1.15. Different views of prototype made by group in ME696 39

Fig. 2.1. Construction of the Prototype Socket 43

Fig. 2.2. Kistler 9272 Dynamometer 45

Fig. 2.3. Construction of Prototype Liner 48

Fig. 3.1. Attachment of “peg” to Socket 52

Fig. 3.2. Instron Testing Set-Up 53

xi

Fig. 3.3. Time vs Load Plot from Instron Testing 53

Fig. 4.1. Feedback Loop of Electrical Components 57

Fig. 4.2. Tekscan FlexiForce Sensor -Model A201 58

Fig. 4.3. Components of Tekscan FlexiForce Sensor 59

Fig. 4.4. Arduino Model ATmega2560 61

Fig. 4.5. Circuitry of Force Sensors and Arduino 62

Fig. 4.6. Rover 5 Motor Driver Board 63

Fig. 4.7. Circuitry of Force Sensor, Arduino, and the Motor Driver 64

Fig. 4.8. Arduino Code 64

Fig. 4.9. Panoramic View of Circuitry Testing Set-Up 66

Fig. 4.10. Arduino Output Results 68

Fig. 4.11. Arduino Output: Close-Up View at Low Force 68

Fig. 4.12. Arduino Output: Close-Up View at an Increasing Force 69

Fig. 4.13. Motor Driver Output Results 70

Fig. 4.14. Motor Driver Output: Close-Up View at Low Force 70

Fig. 4.15. Motor Driver Output: Close-Up View at Increasing Force 71

Fig. 5.1. Initial Set-Up of Prototype Components (Sensors Inserted) 72

Fig. 5.2. Initial Set-Up of Prototype Components (Sensors Not Inserted) 73

Chapter 1: INTRODUCTION

Human locomotion involves the transformation of a series of controlled and

coordinated angular motions occurring simultaneously at the various joints of the lower

extremity into a smooth path of motion. According to Radcliffe (1962), there are six

major factors that influence the path of motion: interaction at the knee and ankle, knee

flexion, hip flexion, pelvic rotation about the vertical axis, lateral tilting of the pelvis, and

the lateral displacement of the pelvis.

The gait cycle may be broken into two parts – the stance phase and the swing phase.

For any given leg, the stance phase begins when the heel makes contact with the ground

and ends at the toe-off position where the foot loses contact with the ground. The swing

phase begins at the toe-off position and ends at heel contact. The major function of the

knee, ankle, and foot during the heel-contact phase is smooth absorption of the shock of

heel contact and maintenance of a smooth path of the center of gravity of the whole body.

The overall objective in the swing phase is to get the foot from one position to the next in

a smooth manner while clearing obstacles of terrain (Radcliffe, 1962).

When an amputation of a lower limb occurs, the individual does not have a “support”

that can absorb the shock when walking. The soft tissues of the limb are forced to absorb

these forces instead. However, because the soft tissues in the limb are not adapted to

great amounts of force, injuries and discomfort can occur. In order to protect these

structures on the amputated side, it is necessary for the prosthetic socket to maintain the

forces and moments within safe limits. As the socket is worn over a certain amount of

time, the forces can compress the limb and cause it (the limb) to reduce in size, thus

making the socket loose and uncomfortable.

2

The objectives of this research are to prove an electromagnetic system is capable of

prosthetic suspension and to quantify this capability. A prototype was built with a force

sensor feedback control loop for the electromagnets, which serves as the foundation for

this study. This thesis focuses on the design and testing of the prototype.

1.1 Amputation: An Overview

Each year, an estimated 158,000 persons are admitted to a hospital nationwide to

undergo an amputation procedure at an estimated direct cost of $4 billion. Many of these

people are either elderly, have dysvascular issues, and/or have diabetes. Although

amputation due to dysvascular causes represents a large percentage of overall

amputations, amputation may also be due to trauma and malignancy (Pezzin, 2004).

According to the1996 National Health Interview Survey (NHIS), it was estimated that

approximately 1.2 million people were living with the loss of a limb. By 2005, it was

estimated by Ziegler-Graham et al. (2008) that there 1.6 million people living with the

loss of a limb in the United States. Table 1.1 illustrates how prevalence varies by etiology

of the limb loss, age, sex, and race. Of the 1.6 million people living with the loss of a

limb, amputations secondary to dyvascular disease (n = 846,000) account for most cases

(54%), and of these, over two thirds have a diagnosis of diabetes (n = 592,000, 70%).

Limb loss due to trauma accounts for an additional 45% of prevalent cases (n = 704,000).

The remaining less than 2% (n = 18,000) accounts for amputations due to cancer.

It must be noted that these percentages vary by age because of the patterns in the

incidence of the underlying disease or injury resulting in amputation. Over two thirds of

amputations due to trauma occur among adolescents and adults below the age of 45

years. Across all categories, 42% of the persons living with the loss of a limb are 65 years

3

or older (n = 665,000), 65% are male (n = 1,026,000) and 42% are non-white (n =

652,000). A total of 65% of people living with the loss of a limb received an amputation

to a lower extremity (n = 1,027,000) and over one half of these amputations were major

(n = 623,000). Of the total number living with the loss of an upper limb, only 8% (n =

41,000) were categorized as major. Estimates of prevalence by type and level can be seen

in Table 1.2 (Ziegler-Graham et al., 2008).

Table 1.1: Age-Specific Estimates of Prevalence by Sex, Race, and Ethnicity, and

Etiology (in thousands): Year 2005, United States (Ziegler-Graham et al., 2008)

4

Table 1.2: Estimates of Prevalence by Type and Level of Limb Loss and Etiology (in

thousands): Year 2005, United States (Ziegler-Graham et al., 2008)

According to Ziegler-Graham et al. (2008), the number of people living with the

loss of a limb is expected to more than double from 1.6 million in 2005 to 3.6 million in

2050. It is predicted that amputations due to dysvascular diseases will account for most of

the increase, increasing from less than 1 million in 2005 to 2.3 million in 2050. The main

cause in increase is due to the aging population and the high rates of dysvascular disease

among older adults. The prevalence of diabetes in the United States is predicted to nearly

double by the year 2030, and nearly triple by the year 2050, solely because of changes in

the demographic composition of the population. Also, given the increase of prevalence of

obesity and the known relationship between obesity and diabetes, a projected increase in

the incidence of amputation due to obesity is likely (Ziegler-Graham et al., 2008)

Ziegler-Graham et al. (2008) also found that limb loss can also be related to race

and ethnicity. Among the underrepresented minority populations in the United States, the

risk of amputation has been noted to be two to three times that of non-Hispanic whites.

Recent studies have been suggested that this variation in risk by race and ethnicity may

be due to poverty and differences in access to primary care and preventative services.

5

One in 250 white Americans are living with the loss of a limb, compared with one in 90

nonwhites. Table 1.3 shows the projected prevalence of limb loss in the United States

over the years until 2050. Amplification of the increase in prevalence of limb loss

associated with dysvascular disease is displayed in Fig. 1.1, assuming ± 10% - 25% in

incidence (Ziegler-Graham et al., 2008).

Table 1.3: Projected Prevalence of Limb Loss by Etiology and Age (in thousands): Years

2005, 2010, 2020, 2050, United States (Ziegler-Graham et al., 2008)

6

Fig. 1.1. Projected Number of Americans (1000s) Living with Limb Amputation

Secondary to Dysvascular Disease (Based on a Hypothetical Change in Incidence on

Amputation post-2005) (Ziegler-Graham et al., 2008)

1.2 Biomechanics of Lower Extremity Sockets

The first attempt to describe force patterns at the limb/socket interface was by

Radcliffe in 1962. Radcliffe proposed a force distribution pattern at the limb/socket

interface which changes throughout the gait cycle. This pattern was influenced by the

alignment of the prosthesis, muscle action, and the angular position of the stump with

respect to the ground reaction force (GRF). Radcliffe assumed that the amputee can walk

in a manner similar to that of an able-bodied person by compensating for the loss of the

ankle function with hip and knee action (Laing et al., 2011).

Three phases of gait were considered: heel strike, mid stance (or shock absorption),

and toe-off. As the heel strikes, the hamstrings prevent the GRF’s acting anterior to the

knee center, causing the knee to extend. Within the socket, the action of the hamstrings

causes high pressure at the patellar tendon and in the posterior distal tibia. Immediately

7

following heel-strike, the GRF passes from a location posterior to the knee joint to a

position anterior to the joint. According to Radcliffe (1962), this results in the largest

change in pressure as the knee extension moment changes to flexion (Laing et al., 2011).

During mid-stance (from flat-foot to heel-off), the knee undergoes controlled flexion.

Though Radcliffe proposed the largest change occurs immediately after heel strike, a

study by Goh et al. (2003) found that the interface pressure showed the greatest change

between mid- and late stance. Contributing to these differences were factors outside of

the GRF’s that were not considered by Radcliffe, such as alignment, thigh muscle

strength, and stump morphology. As the body advances over the stabilized knee, the

GRF’s act posterior to the knee. Action by the quadriceps and forceful extension of the

hip prevents the knee from collapsing. As a result, forces are concentrated at the patellar

tendon, anterior distal tibial and popliteal area (Laing et al., 2011).

During toe-off, the GRF passes behind the knee and, as such, the same three areas

experience high pressure. In addition to the anterior-posterior forces are the medial-lateral

forces, which, as Radcliffe proposed, are relatively similar throughout the gait cycle.

GRF’s have a medial inclination due to the horizontal acceleration of the center of

gravity. As a means to counteract this medial inertia, stabilizing forces are established in

the lateral distal and proximal medial tibia (Laing et al., 2011).

1.2.1 Shear, Slippage, and Friction

The pressures at the liner and socket interface vary among sites, individuals, and

clinical conditions. The basic principles for socket design vary from either distributing

most of the load over specific load-bearing areas or by distributing the load more

uniformly over the entire limb. Still, no matter what kind of design, it is always important

8

to factor in the load-transfer pattern. This will help designers to evaluate the quality of

fitting and enhance their understanding of the underlying biomechanical rationale (Mak

et al., 2001).

The skin and underlying tissues of the residual limb are not particularly adapted to

the high pressures, shear stress, abrasive relative motions, and the other physical

irritations encountered at the socket interface. The pressure distribution between the limb

and socket is a critical consideration when it comes to socket design and fit. It is critical

to understand how the residual limb tissues respond to the external loads and other

external phenomena at this interface (Mak et al., 2001).

1.2.1.1 Tissue Responses to Mechanical Loading: Normal and Shear Force

Tissue responses to external forces include many things such as tissue

deformation, interstitial fluid flow, ischemia, reactive hyperemia, sweat, pain, skin

temperature, skin color, and more. An application of either an unusually very large force

or a prolonged or repetitive force may damage function and/or structures. Mechanically,

forces applied to the skin surface will produce stresses and strain within the skin and

underlying tissues. Those stresses affect cellular functions and other processes in the

tissues. When moderate static forces are applied to the skin, the underlying blood vessels

and lymphatic drainage can be obstructed or partially obstructed, and oxygen and other

nutrients can no longer be delivered at a sufficient rate that satisfies the metabolic

requirements of the tissues. Without proper circulation, the breakdown products of

metabolism would gather within the tissues and affect the cellular function. Although a

moderate force may not cause direct and immediate damage to the tissues, a repetitive

forces may. Repeated applications day after day could cause an inflammation reaction

9

and may even result in tissue necrosis. If the applied load is within a certain range, tissue

adaption may occur by changing its tissue composition and architecture (Mak et al.,

2001).

Besides the magnitude of the load, other load characteristics (such as direction,

distribution, duration, and loading rate) should also be considered when fitting a

prosthetic socket. Two types of localized stresses are generated between a residual limb

and a prosthetic socket during ambulation: normal stresses (perpendicular to the

interface) and shear stresses (in the plane of the interface), as shown in Fig. 1.2. Normal

and shear stresses are important because they can traumatize limb tissues (Sanders et al.,

1993). When pressure are evenly distributed over a wide area of the body, damage is less

than if the load were applied over a localized area. An inverse relationship exists between

the intensity of the external loads and the duration of the load application (Mak et al.,

2001).

Fig. 1.2. Diagram of Normal and Shear Stress Applied at a Certain Point on the

Socket (Sanders and Daly, 1993)

Residual limb soft tissues within a prosthetic socket are subjected to many things.

First, pressure and shear forces are applied by the socket when it is fitted on the limb. The

10

soft tissues of the limb are not suited for undertaking such dynamic and repetive loads

during ambulation. Second, skin rubbing against the socket edge and interior surface may

happen. This will result in sporadic skin deformation and biomechanical irritations. If

excessive slippage occurs between the skin and socket, tissue abrasians can occur and

heat will be generated. Third, residual limb tissues will be often exposed to high-

humidity enviromnents. This is due to the socket being fitted intimately on the limb,

eliminating the ability of air circulation and trapping accumulated sweat. Fourth, the

residual limb may suffer from possible chemical and mechanical irritations or allergic

reactions to various socket materials. Under such conditions, the residua limb soft tissues

may either adapt or break down (Mak et al., 2001).

1.2.1.2 Slippage

The biomechanics of the coupling between the limb and socket is an important

factor for socket fit. This coupling is affected by the relative slippage between the

subject’s skin and the prosthetic socket, and by the changes of the residual limb tissues.

The coupling stiffness could be influenced by the tightness of the fit. The tightness of the

fit can be altered by the change in pressure distribution caused by the variations in socket

shape. Generally, a loose fit may allow for more slippage, which may affect stability,

while a very tight fit may cause a more stable connection yet increase the interface

pressures (Mak et al., 2001).

Another important factor affecting slippage is the friction between the subject’s

skin and prosthetic. Extreme amounts of slippage at the socket interface should be

avoided, however, absence of slippage may cause other issues as well. Discomfort may

11

arise, not from the pressure, but from the increase in interface temperature and perspiratin

inside the socket (Mak et al., 2001).

1.2.1.3 Frictional Forces

Friction is the phenomenon in which tangential forces are acting between bodies

in contact oppose their relative motion or impendinding motion. Friction between the

residual limb and socket leads to two primary effects. The first being that friction

produces shear forces on the skin, leading to tissue distortion. This may disrupt tissue

functions and can be harmful to the tissues (Mak et al., 2001).

The second effect is that the friction-producing shear forces at the skin surface

can assist in supporting the ambulatory load and help with the supension of the prosthesis

during the swing phase. A study by Zhang et al. revealed that the smaller amount of

friction leads to a smaller amount of shear stress. This will increase the amount of normal

stresses required to support the same load. Hence, reduction of interface friction may not

always be a good way to solve residual limb tissue problems. An adequate coefficient of

friction could help to support loads and prevent unwanted slippage. However, a surface

with large frictional forces could experience high local stresses and tissue distortion when

donning the limb into the socket and during ambulation. A balance between the

requirements for effective prosthetic control and minimization of interfacial harards could

be achieved when suitable amounts of frictional forces are applied (Mak et al., 2001).

Research related to friction in the socket has been done and includes 1)

investigating the coefficient of friction of skin with various interface materials, 2)

measuring the amount of shear stresses and slippage at the interface, 3) measuring the

amount of relative motion between the limb and socket, and 4) contributions of frictional

12

shear to the load transfer. Frictional properties of human skin have also been investigated

under various skin conditions to examine the effects of skin care products and to see how

friction might affect some friction-dependent manual activities.The skeletal movements

relative to the socket are determined by the relative slip between the skin and the

prosthetic socket, as well as by the deformation of the residual limb tissues (Mak et al.,

2001).

1.3 Interstitial Fluid

In the human body and circulatory system, there are many cells. These cells are

linked to each other with the help of intercellular connections. In between these cells is a

fluid that bathes the cells – interstitial fluid. Interstitial fluid is the main component of

extracellular fluid, which is found in interstitial tissue spaces. The basic composition of

interstitial fluid is water, along with solutes like sugars, fatty acids, amino acids, salts,

urea, white blood cells, etc. Although the basic composition of the fluid remains the

same, it differs slightly depending on the region of the body where the fluid is present

(Interstitial Fluid, 2010).

Interstitial fluid is responsible for basically maintaining the homeostasis in the

cell and in the body. This is the fluid that helps to deliver nutrients to the cells and also

helps to carry waste from the cells for elimination. When a certain disease is located in

the body, interstitial fluid removes it from the cells and puts it back into circulation. If the

fluid is not removed, then there is the possibility for a buildup of this fluid, which leads to

swelling (Interstitial Fluid, 2010).

13

Fig. 1.3. Interstitial Fluid (OpenStax College, 2013)

1.3.1 Limb Shrinkage

In general terms, the issue is the balance between the hydrostatic and osmotic forces

within the blood vessels (intravascular pressures) and those in the surrounding tissue

(interstitial pressures). Within the vessel the primary hydrostatic force is the force of the

heart as it pumps blood when it contracts (systole). This is normally stronger than any

hydrostatic force acting on the interstitial space so the tendency is for fluid to have net

movement out of the vessels due to hydrostatic pressure. Colloid osmotic (oncotic)

pressure, is the force that tends to draw fluids back into the vessel. Oncotic pressure is

created by proteins that normally cannot leak out through the vessel membrane and thus

create a consistent pressure promoting fluid movement into the vessels.

Fig. 1.4. Pressure Distribution between Blood Vessel and Interstitial Fluid (Marieb,

2003)

14

Increased pressure on the soft tissue will be transmitted into the interstitial spaces,

resulting in an increase in the hydrostatic pressure of the interstitial spaces. This will

force fluids which are “trapped” locally in the interstitial space to move into the venous

and lymphatic systems. Both of these systems have one-way valves and serve to return

fluids to the central circulation where they are recirculated by the heart. This fluid

transfer can occur at the level of the capillaries. Capillaries are neither arteries nor veins,

but the intersection between the two. Fluids that leave the vascular system from the

arterial side (pumped away from the heart) can only re-enter on the venous side

(recirculated back to the heart).The capillaries form web-like networks within the muscle

and soft tissue called capillary beds. Within the capillary beds are the lymphatic

capillaries. The lymphatic system also serves to return fluids to the central circulation.

This movement of fluid from the limb to the central circulation causes a reduction in

volume of the limb. So long as the prosthesis applies pressure to the soft tissue, the

hydrostatic pressure within the interstitial spaces of the limb will remain elevated, so that

fluids removed from the limb will not return. At some point an equilibrium is reached

where the tissue cannot be compressed further. In amputees these forces are deranged by

multiple factors. Illness and trauma can damage blood vessels, increasing their

permeability. Chronic illness can lead to malnutrition which decreases blood protein and

reduces oncotic pressure within the vessel. Both factors can lead to increased interstitial

fluid. The use of a prosthetic device, with weight bearing on surfaces that were not

intended to be weight bearing, increases the hydrostatic pressure in the interstitial space.

If hydrostatic pressure is greater and osmotic pressure is lower in the interstitial space

15

than the pressures of the intravascular space there will be a net movement of fluid into the

vessel.

All the above results in a pattern commonly seen in lower extremity amputees in

which the limb swells when not in the prosthesis, then shrinks when the prosthesis is

worn. Vacuum-assisted suspensions and our proposed design provide what in essence is a

negative pressure environment to reduce the interstitial hydrostatic pressure. Wearing the

prosthetic device for long periods of time may cause the residual limb to shrink. A

common complication of prosthetic device use is residual limb volume fluctuations

during the use of the prosthesis (Gerschutz et al., 2010). Many amputees are hindered by

the necessity of constantly managing residual limb volume with the application of socks

and volume management pads. This is due in part to pressure from the prosthetic socket

on the limb compressing tissue that is not meant to support such high levels of weight.

Other medical factors, including the user’s internal fluid volume status, may also have an

effect.

The stabilization of residual limb volume is important for comfort, reduce tissue

breakdown, and improve daily function. In a recent study, residual limb volumes

fluctuated as much as 6.5% over the course of a 24 hour period (Gerschutz et al., 2010).

Loss of residual limb volume can lead to loosening, improper distribution of weight

bearing forces, and poor fit of the limb in the socket. This can cause increased pressure

and shear force on the limb, which can potentially lead to skin breakdown and a loss of

control over the prosthesis.

16

1.3.2 Physiology of Residual Limb Shrinkage

Positive pressures decrease the volume of the limb while negative pressures increase

limb volume. Vacuum suspension further improves proprioception and control of the

prosthesis by preventing the limb from losing volume during the day. Unlike other

designs of suspension where the socket fit becomes sloppy as the limb loses volume each

day, the limb stays hydrated and set to the socket (Street, 2007).

Limb volume fluctuates as pressure fluctuates. In the morning before putting the

socket on, with one atmosphere of pressure (1atm), limb volume is stable. As the limb

pressure increases, for example after placing on an undersized socket or during stance,

the limb loses volume. This is because elevated pressure (pressure greater than 1atm)

increases the amount of interstitial fluid being driven back into the bloodstream and

lymphatic vessels, and out of the limb. However, in contrast, if the pressure drops below

1atm, such as when the tibia extracts and causes the soft tissues to elongate during the

swing phase, the limb gains volume. This is because low pressure (pressure less than

1atm) increases the amount of fluid being drawn out of the blood stream and into the

limb’s tissues (Street, 2007).

Sustaining proper circulation and fluid exchange in the soft tissue is imperative for

maintaining a healthy residual limb. Pressure applied to the limb by the socket system,

chiefly during ambulation, complicate this task. Therefore, clinicians need to be aware of

the pressure applied to the residual limb when prescribing socket systems to patient (Beil

and Street, 2004).

17

1.4 Prosthetic Devices

There are many different types of prosthetic devices. Prosthetics systems are available

for majority of the body and come in an array of designs. Different types of prosthetic

limbs are designed with different goals in mind. Often, these goals depend on the site of

amputation and the needs of the patient. For example, a cosmetic prosthetic limb (called a

cosmesis) is designed with appearance in mind rather than controllability. Advanced

plastics and pigments are matched to the patient’s skin tone and allow for a more lifelike

appearance. On other hand, other prosthetic limbs are designed with usability and

function as a central purpose. An example of this would be a common controllable

prosthetic hand that might consist of a pincer-like split hook that can be opened or closed

to grip objects or perform other tasks. Body-powered and externally powered prosthetic

limbs are also available, however, the price of prosthetic limbs can be very high. An

“upper extremity” amputation involves the loss of all or part of an arm (sometimes both).

A “lower extremity” amputation involves the loss of a certain portion of one or both legs

(Clements, 2008).

1.4.1 History of Prosthetics

A prosthesis or artificial limb is a device that aim to substitute the loss of a limb with

cosmetic and functional desirability for the amputee (Laing et al., 2011). Acient literature

contatins references to prosthetic limbs in stories and poems, but some of the earliest

historical accounts of prosthetic limb use were recorded in Greek and Roman times. In

the year 2000, researchers in Cairo, Egypt discovered what they believed to be the oldest

documented artificial body part – a prosthetic toe made of wood and leather. This is

shown in Fig. 1.5. The device, found attached to the remains of a 3,000 year old

18

mumified Egyptian noblewoman, dates back to between 950 and 710 B.C. (Clements,

2008).

This device represents how little prosthetic limb design has changed throughout

history. Nearly 2,000 years later during the Dark Ages, armored knights often relied on

iron prosthetic limbs that were usually crafted by the same metalworker that created their

armor. These limbs were bulky and not very functional, yet were used more for the

purpose of hiding the limb loss (Clements, 2008).

Most famously attributed to pirates, peglegs with wooden cores and metal hands

shaped into hooks were the prosthetic standard throughout much of history. While

Hollywood may have exaggerated their use of hook and peglegs, pirates did rely on these

types of protheses. The materials for these could be scavenged from a common pirate

ship and the ship’s cook typically performed the amputation surgeries (Clements, 2008).

In the early part of the 16th centery, a French militrary doctor named Ambroise Paré

invented a hinged mechanical hand as well as prosthetic legs that featured advances such

as locking knees and specialized attachment harnesses. Around 1690, a Dutch surgeon

named Pieter Verduyn later developed a lower leg prosthesis with specialized hinges and

a leather cuff for improved attachment to the body (Clements, 2008).

As artificial limbs become more common, advances in areas such as joint technology

and suction-based attachment methods continued to advance. In 1812, a prosthetic arm

was developed that could be controlled by the opposite shoulder with connecting straps.

In 1945, the National Academy of Sciences, an American government agency,

established the “Artificial Limb Program”. This program was created in response to the

influx of World War II veteran amputees and for the purpose of advancing scientific

19

progress in artificial limb development. Since this time, advances in areas such as

materials, computer design methods, and surgical techniques have helped prosthetic

limbs to become increasingly lifelike and functional (Clements, 2008).

Fig. 1.5. Prosthetic Toe Found in Egypt (Clements, 2008)

1.4.2 Modern Prosthetic Limbs

Compared to those of historical times, one major difference in modern prosthetic

limbs in the presence of newer materials such as advanced plastics and carbon-fiber

composites. These materials can make a prosthetic limb lighter, stronger, and more

realistic. Electronic technologies help control the prosthetic, allowing it to automatically

adapt its function during certain tasks. Still, while new materials and technologies have

modernized prosthetics, the basic components of the prosthetic remains the same

(Clements, 2008).

Lower extremity prostheses consist of an assembly of several components such as the

shank (also known as the “pylon”), socket, ankle, and foot (Fig. 1.6) (Laing et al., 2011).

The shank is the internal frame of the prosthetic limb. The shank must provide structural

20

support and has been traditionally formed of metal rods. In more recent times, lighter

carbon-fiber composites have been used. They are sometimes enclosed by a cover,

typically a foam-like material, and the cover can be shaped and colored to match the

recipient’s skin tone to give the prosthetic a more lifelike appearance (Clements, 2008).

The socket is the portion of the prosthetic that interfaces with the patient’s residual

limb. Because the socket transmits forces to the patient’s body, it must be fitted exactly to

the residual limb to ensure that it does not cause irritation or damage. Current prosthetic

devices utilize a liner and socket to attach to the limb (Clements, 2008).

Still, with any prosthesis, the body needs time to adapt. Not only is a poor suspension

system inadequate, but an uncomfortable socket fit is a common complaint from lower

limb amputees with surveys revealing that amputees believe comfort is the most

important aspect of the prosthesis. These surveys also revealed that over half of all

wearers are in moderate to severe pain for most of the time while wearing the prosthesis

(Laing et al., 2011). Pain and discomfort may be caused by either the suction and/or the

socket type.

Fig. 1.6. Typical Trans-Tibial Prosthetic (Laing, 2011)

21

1.4.2.1 Types of Prosthetic Suspension Systems

The suspension system for attaching the prosthetic limb to the body is critically

important in prosthesis. The suspension system of a prosthetic device is what holds the

device onto the residual limb. There are many types of prosthetic suspension systems.

The simplest type of suspension method may include a harness system, straps, or belts.

There are more advanced suspension methods available, but each have their own

advantages and disadvantages. However, none of these systems fully satisfy the needs of

the user when it comes to prosthetic control and control of limb volume changes.

1.4.2.1.1 Sleeve Suspension

In sleeve suspension, a material with a high coefficient of friction is placed over

the socket and limb. With this high coefficient of friction, the sleeve does not slide and

hold in everything in place. This type of socket uses the principle of surface tension to

hold the socket on and must be put on the limb before the prosthesis is put on.

Still, the problem with sleeve suspensions is that they can easily rip or tear, causing the

seal to break. Also, it is possible for the sleeve to become loose due to such things as oils

and sweat. For some users, the sleeve may act as another layer needed to put on and may

feel uncomfortable. The sleeve suspension also reduces the range of motion of the limb

as the material may bunch behind the knee. Being as it is only a material that covers the

prosthesis and limb, it does not provide any support towards limb volume fluctuations.

22

Fig. 1.7. Sleeve Suspension (Cornell, 2011)

1.4.2.1.2 Pin and Lock Suspension

The pin and lock mechanism are easy to use and frequently recommended for an

amputee’s first prosthetic. Pin and lock systems use gel liners within undersized total

surface-bearing (TSB) sockets. Pin suspension uses a metal pin extending distally from

the liner that locks into a receptacle at the bottom of the socket (Beil and Street, 2004).

The pin and lock system provides a reliable lock that allows for easy attachment and

removal.

However, this type of suspension is incapable of dealing with torsional forces and

results in rotation of the prosthetic. Some symptoms most commonly seen in amputees

who use a locking pin system for suspension are daily reddening and swelling of the

distal residual limb. Long-term changes with pin use include general thickening and

discoloration of the distal tissues (Beil and Street, 2004).

In addition, the most important disadvantage with this system is that the pin and lock

suspension does not have any mechanism to help with differences in limb size. With pin

and lock suspension, during the swing phase, there is a very large pressure drop. This

23

paradox is due to the fact that, while pin/lock suspension is even more forceful in

drawing fluids into the limb, it is only at the distal end of the limb. The proximal portion

of the limb is squeezed. So with the pin/lock suspension, instead of moderate to strong

global filling of the limb, there is only strong distal filling with a tendency for congestion

of the fluids and volume loss proximally (because of the simultaneous proximal squeeze)

(Street, 2007).

Fig. 1.8. Pin and Lock Suspension (Coalition, 2008)

1.4.2.1.3 Suction Suspension

The idea for suspending a socket through suction was patented in the United

States (US) in 1863, but was not utilized frequently in the US until World War II when

US military surgeons and engineers noted the success that Germany was having with the

suction suspension valve on their veteran amputees. Today, suction is the primary means

of socket suspension ultilized by the military amputee population. (Harvey et al., 2012)

Like the sleeve, a material with a high coefficient is placed over the limb and

prosthesis. In this type of suspension, suction develops in the slight air space between the

gel liner and socket when the liner attempts to slide proximally relative to the socket

24

during the swing phase of ambulation. The air space is sealed by a gel sleeve covering the

proximal socket and liner along with a one-way valve at the distal expulsion port of the

socket (Beil and Street, 2004). Still, as fluid is moved out of the limb into circulation due

to the pressure, suction does not have a mechanism to draw the fluid back in. As a result,

suction only cause the limb to reduce in size.

Fig. 1.9. Suction Suspension (Coleman, 2004)

1.4.2.1.4 Vacuum Suspension

The current “state of the art” technology for prosthetic attachment utilizes vacuums

(Itoga et al., 2013). In a vacuum suspension, a vacuum pump removes air molecules from

the thin, sealed air space (sheath) between the total surface bearing (TSB) socket and

liner, as shown in Fig. 1.11 (a) The vacuum created by the removal of the air molecules

holds the liner firmly in place and globally to the walls of the socket, as shown in Fig.

1.11 (b). The vacuum applies negative pressure on the residual limb. The limb is

completely isolated from the vacuum. The principle is that the liner, and therefore the

skin, are no longer able to separate from the socket. Eliminating the separation between

25

the liner and socket improves the patient’s spatial awareness and control over the

prosthetic device (Street, 2007). This also maintains an effective intimate fit, which deals

with torsional forces exceedingly well.

This lack of separation of the liner and limb from the socket is thought to explain why

the vacuum suspension prevents volume loss and improves limb health. A nearly

universal observation with vacuum suspension users is the reduction or elimination of

minor skin problems. In some cases, open wounds healed and remained healed upon

switching to vacuum suspension. The first trans-tibial amputees to use vacuum

suspension in 1999 also reported that their limbs no longer lost volume during the day

(Street, 2007).

Still, the benefits of vacuum suspension are only recognized by the amputee if the

limb and liner are in total contact with the socket and an air tight seal is created. The

outer sleeve used to create the vacuum seal is large and cumbersome and can wear out

relatively quickly. Users have also complained that the vacuum seal can break if they sit

for extended periods of time. Systems that use vacuum apply negative pressure to the

entire socket and there is no capability to adjust the pressure in different locations.

The prosthetist must design and construct a TSB socket that closely matches the

shape of the amputee’s limb and is free of specific weight bearing structures and areas of

relief. If both the prosthetist and amputee meet this requirement and vacuum is

maintained, the amputee will realize the benefits of vacuum suspension. If they fail to

meet this requirement, the liner/limb will become detached from the socket and the limb

will experience pressure and skin damage in extreme cases (Street, 2007).

26

Fig. 1.10. Vacuum Suspension (Hanger, 2014)

(a) (b)

Fig. 1.11. (a) Cross section of vacuum suspension system showing sealed air space. (b)

The vacuum creates forces that anchor the liner to the socket. The sum of all the axial

components of the axial forces creates a large suspension force (Street, 2007).

27

1.4.2.2 Types of Prosthetic Sockets

Besides prosthetic suspension systems, pain and discomfort due to limb changes may

also be due to the type of prosthetic socket. The socket is considered as the most

important aspect of the artificial limb and constitutes the critical interface between the

amputee’s residual limb and prosthesis. The design and fitting of the socket is considered

as the most difficult procedure due to the uniqueness of each amputee’s stump. Still,

socket designs have evolved from the basic conical shape to much more advanced

designs.

1.4.2.2.1 Patella Tendon Bearing Socket

The end of World War II saw the discovery of new materials and a greater

understanding of biomechanics. As a result, it started the beginning of a more

modernized approach to socket design. Prior to these designs, thigh corsets were often

utilized to off-load the stump and sockets were loaded only around the proximal edges.

However, this design allowed for the movement of the limb in the socket, which resulted

in skin irritations and pain (Laing et al., 2011).

In 1957, the patella tendon bearing (PTB) socket was introduced for below-knee (BK)

amputee patients at the University of California at Berkeley (Hachisuka et al., 1998). This

design was the first to remove the corset and sidebars so that the entire weight of the

amputee was taken at the stump/socket interface. The socket takes advantage of the

pressure tolerant areas in the stump, especially that of the patellar tendon and the

posterior aspect of the stump. As such, the socket shape is indented to increase the load

on areas that are more pressure tolerant, commonly known as the patellar tendon bar (Fig.

28

1.12). The patellar tendon bar helps to relieve loading at the other regions of the stump

that are considered less tolerable to load, reducing discomfort.

However, considerable skill is required in order to generate a good PTB socket fit.

Also, because each residual limb is different, this type of socket may can actually cause

pain and discomfort for some. Patients wearing the PTB socket may have complaints

about excessive pressure on the patellar tendon area, limitation of knee flexion,

adventitious bursae, skin abrasions, or dermatitis. These problems may result from the

socket design (because the socket does not suspend the prosthesis from the stump, it

needs supplementary suspension) and it cannot prevent pistoning and sliding on the skin

(Hachisuka et al., 1998).

Fig. 1.12. Patellar Tendon Bearing Socket (Laing, 2011)

1.4.2.2.2 Total Surface Bearing Socket

To solve the problems of the PTB socket, another type of socket was created - the

total surface bearing (TSB) socket. Advantages in suction suspension in the 1950s led to

the development of a total-contact silicone gel-lined socket (Hachisuka et al., 1998).

Using suction as the suspension method, this socket is a total contact suction socket in

which weight was born on the entire surface of the stump, including areas previously

considered pressure sensitive (Hachisuka et al., 1998). The uniform pressure causes an

elongation of the soft tissue surrounding the stump, resulting in more padding at the

29

sensitive distal end and firmer tissue consistency. The TSB socket can provide an

increased range of motion, uniform pressure, and decreased weight perception (most

likely to the increased suction) (Laing et al., 2011).

One example of the TSB concept is the Icelandic Roll-On Silicon Socket (ICEROSS)

developed in the mid-1980s by Össur Kristinsson. Using a silicon liner turned inside-out

and rolled over the residual limb, the liner forces the limb’s skin in a distal direction,

stabilizing the soft tissue. In addition to the liner, padding is placed over the bony areas of

the limb during the casting process (Laing et al., 2011). The ICEROSS came in six sizes

and was rolled over the stump as an inner socket and suspension component to provide

good total contact with the limb skin (Hachisuka et al., 1998)

Because this socket was in contact with the entire stump surface, it was thought that

the improved pressure seal provided sufficient suction suspension and that the dispersion

of local pressures was more comfortable. The TSB socket provides less piston movement

during walking, less traumatization of the skin of the limb, and less interference with

knee flexion than experienced with the PTB socket. However, some subjects using the

TSB socket may complain about pain at the distal end of the limb, discomfort during

knee flexion, perspiration and odor of the limb, and problems with donning and doffing

the socket. Other less frequent problems were a periodic feeling of tightness around the

mid-portion of the limb and unravelling of the proximal end of the silicone inner socket

(Hachisuka et al., 1998).

1.4.3 Leading Prosthetic Suspension Companies

The major competitors providing vacuum suspension systems are The Ohio

Willow Wood Company ($30M revenue, 160 employees) and Otto Bock Healthcare

30

($906M revenue, 5,196 employees), with the LimbLogic and Triton Harmony

technologies, respectively. Endolite, another major company, produces mainly pin & lock

devices. Two companies that specialize in liner and sleeve production are ALPS and

Iceross. Heavy wear on liners allows these companies operate on a disposable business

model. Individuals that live very active life styles will go though one inner liner per week

and in the case of vacuum systems, outer sleeves must be replaced once every six months

(Itoga et al., 2013).

There are a few other technologies that exist in the market but have not gained

wide acceptance. Peak Prosthetic Designs produces a prosthetic that integrates a BOA

ratcheting device to allow the wearer to adjust the fit of the socket. Another company, CJ

Socket Technologies, tried to solve the problem of fit degradation by redesigning the

socket entirely.

The CJ Socket uses a hard J-shaped plate in combination with a flexible non-

elastic “sail” to create a fit that can be adjusted with the Velcro straps attached to the sail.

While both of these solutions allow users to adjust the fit throughout the day, neither of

them does anything to prevent limb volume loss. Additionally, neither the CJ Socket nor

the Peak Prosthetics Designs’ socket addresses the effects of torsional forces affecting the

alignment of the prosthetic. While this is less of a problem with the CJ Socket than pin

and lock sockets, it does remain unaddressed by the design. Though these mechanical

systems do not often fail, they are typically seen as suboptimal for an amputee with an

active lifestyle, as the problems with torsion pose a danger to an active prosthetic user

(Itoga et al., 2013).

31

1.5 Literature Review

This section presents the foundation of the literature found for this research. The

journal articles related to residual limb shrinkage are described. The experimental

literature review presents experimental studies related to the comparison of different

socket and suspension types.

1.5.1 Residual Limb Volume

Street (2007) found that, as pressure on the limb increases, the limb loses volume.

Volume is lost because elevated pressure (greater than 1 atm) increases the amount of

interstitial fluid being driven back into the bloodstream and lymphatic vessels, and out of

the limb. Gerschutz et al. (2010) investigated the effects of different vacuum pressure

settings on a lower-limb amputee at three treatment levels: absence of elevated vacuum

(suction), vacuum (negative pressure) at 10 in Hg, and vacuum (negative pressure) at 15

in Hg. The results indicated a significantly less volume fluctuation with vacuum

compared with suction and an improvement in volume retention with length of vacuum

suspension usage.

1.5.2 Comparison of Socket and Suspension Types

Beil et al. (2004) measured limb interface pressures during ambulation with pin/lock

and suction suspension systems. No pressure differences were seen between the two

during stance phase, however, during the swing phase, the pin/lock suspension

maintained a greater average compressive pressure and greater suction force. Yigiter et

al. (2002) investigated the effectiveness of a patellar tendon bearing (PTB) and a total

surface bearing (TSB) socket on prosthetic fitting and rehabilitation. Data analysis

showed that TSB sockets were lighter, better in suspension, and more effective during

32

walking activities than the PTB socket. Sanders et al. (2011) used bioimpedance analysis

to measure the residual limb volume of subjects using elevated vacuum sockets and

nonelevated vacuum sockets. It was found that fluid volume loss were less for elevated

vacuum sockets. Beil et al. (2002) measured interface pressures during ambulation with a

normal TSB suction socket and a vacuum-assisted socket. The vacuum-assisted socket

was shown to create significantly lower positive-pressures during the stance phase and

greater negative-pressures during the swing phase. Ali et al. (2012) investigated the

effects of three dissimilar suspension systems - polyethylene foam liner, silicone liner

with shuttle lock, and seal-in liner. The results showed participants were more satisfied

with the seal-in liner and experienced fewer problems with this liner. Sullivan et at.

(2003) outlined the importance of pre-osseointegration assessment and looked at the

physical, and prosthetic, advantages of direct skeletal attachment. Hachisuka et al. (1998)

investigated the TSB prosthesis for below-knee amputee patients and determined its

clinical indications. It was found that the TSB socket is suitable for and preferred by

many amputee subjects. Klute et al. (2011) investigated the effect of a TSB socket with

vacuum-assist compared with a pin suspension on lower extremity amputees. The TSB

socket with vacuum-assist resulted in a better fitting as measured by limb movement

relative to the prosthetic socket. Coleman et al. (2004) compared two common trans-tibial

socket suspension systems: the Alpha® liner with distal locking pin and the Pe-Lite™

liner with neoprene suspension sleeve. It was found that ten out of the thirteen subjects

preferred the Pe-Lite™ with neoprene suspension sleeve. Board et al. (2001) compared

volume changes associated with normal and vacuum conditions using a TSB suction

socket. With the vacuum system, the limb gained an average of 3.7% in volume.

33

There has not been research published about electromagnetic means of prosthetic

suspension. This is a new idea that becomes the focus of this research.

1.6 Research Motivation

One principle of a lower extremity prosthetic is that, because the residual limb is not

adapted to accepting the forces associated with ambulation, then the force needs to be

distributed across a greater area. The foot is a very bony, tough apparatus, wrapped in

very strong ligaments and tendons and encased in very tough skin. It is adapted by

evolution and daily use to accommodate this force. When this force needs to be

transferred to the soft tissue of the residual limb, this tissue is not adapted to accept that

force. Therefore the prosthetic is formed to transfer the force across a greater surface

area. If the force is too focal, the softer skin and tissues of the residual limb will break

down.

In order to transfer this force in a relatively uniform manner, the prosthetic socket is

formed in roughly a conical shape. Similar to the Roman arch in a 3D form, this

distributes the force of ambulation in a relatively uniform manner. However, most of the

force is transmitted vertically - body weight; gravity; the various forces associated with

stepping up and down, jumping etc. The effect of a large vertical force directed into a

conical shape results in the distal end of the residual limb becoming compressed. The

longer the prosthetic is worn, the more compression occurs. The compression is limited

by the mass of the compressed tissue, the amount of residual limb bone, etc. These forces

of compression also lead to decrease in volume of the residual limb. The compressive

forces push fluid that resides between the cells and tissues (aka the interstitial space) back

34

into the circulation. This causes a progressive decrease in the circumference of the limb,

extending from the distal to the proximal aspect of the limb.

The lateral aspects of the residual limb play a large role in controlling the rotational

and shear forces of the prosthesis. As the general circumference of the limb is reduced,

control over these rotational and shear forces is lost. This makes control of the prosthesis

difficult, particularly when ambulating on uneven terrain. The end result is often referred

to as "pistoning" in which the lateral aspects of the limb no longer contact the prosthetic

socket. In this situation there is almost no control of the prosthesis, and the distal aspect

of the residual limb can be damaged by having to absorb excess weight bearing force.

In a diseased limb, such as is seen with diabetes, the early morning size of the limb

may be enlarged by poor circulation, low serum protein levels etc. However, after even a

short period of ambulation with a prosthesis, these volumes can be significantly

decreased by the forces described above. This can result in poor prosthetic fit leading to

skin breakdown, limitations in activity, all counterproductive to the point of prosthetic

fitting in the first place.

Trying to solve this complicated problem crosses the fields of medicine, prosthetics,

engineering, and materials science. Currently, vacuum suspension is the leading method

to compensate for limb volume changes. However, as mentioned above, the vacuum

suspension has many flaws as well. There is a need for a new prosthetic system with the

capability to adjust and control the pressure at different points in the prosthetic that will

provide a more comfortable fit, more control over the prosthesis, and better means of

curbing limb volume loss. At the moment, a suspension system through the use of

magnets is the best option available – magnets are small and lightweight, and

35

electromagnets have the ability to be “turned on or off”. Though another option could be

to create a socket that is able to change shape in real time, but still stay rigid, the

technology to create this would be very advanced. By combining the best of all fields (i.e.

lighter, holds better, more durable), many amputee could benefit from this.

Fig. 1.13. Comparison between the Electromagnetic Suspension System and the Current

Problem

1.6.1 Contributions by ME 696 Group

This project was first approached by a group of undergraduate students for an ME

696 course. Presented in this section are the liner and socket design concepts that the

undergraduate group created for their prototype called MAGNOLEV. MAGNOLEV is a

revolutionary prosthetic suspension system that utilizes magnets to generate negative

pressure around the residual limb to prevent volume reduction and even promote wound

healing. MAGNOLEV was found to be more durable and more comfortable than the

current vacuum systems. Through research, Itoga et al. found that negative pressures

between 1.955-3.325 kPa are sufficient to prevent residual limb volume loss. The tests

36

validated that magnets can generate sufficient force to achieve this desired pressure (Itoga

et al., 2013).

1.6.1.1 Liner Design

The liner refers to the sleeve that is placed over and is in direct contact with the

residual limb. It is preferential that the liner allow for perspiration evacuation. A

breathable liner would increase comfort as current liners trap perspiration, increasing the

potential for the liner to slip and growth of bacteria. Flexibility will be another

requirement essential for maintaining comfort and facilitate ease donning the liner (Itoga

et al., 2013)

The liner was a standard thermoplastic liner with and outer coating of magnetic

material. The liner was fabricated by rolling a rectangular sheet of thermoplastic that had

been coated with seven coats of magnetic paint and two layers of iron filings material

into a cylinder. The paint coatings and iron fillings allowed the magnets imbedded in the

socket to pull on the liner and create the desired negative pressure. At the bottom of the

liner, a metal plate was embedded in the thermo plastic and left exposed. This was done

as to allow the electromagnet in the bottom of the socket to make direct contact with the

plate. The seams of the thermoplastic were sutured together using needle and thread

(Itoga et al., 2013).

During testing, it was shown that a drastic decrease in holding force as the

distance between the metal plate and the electromagnet increased. Even the liner cloth

diminished the attractive force enough to cause an issue. As such a hole was cut in the

liner cloth to allow for direct contact between the liner and the metal plate. Figure 1.14

shows the liner (Itoga et al., 2013)

37

Fig. 1.14. Liner design from ME696 student team

1.6.1.2 Socket Design

The socket refers to the hard outer portion of the prosthetic that is molded to the

individual’s residual limb. Dr. Roe mentioned that in order to maintain a proper fit and

reduce pressure points; it would be nice to have a system that could dynamically adjust

the forces to alleviate these pressure points. This can be deemed as “adjustable fit”. To

ensure accurate transmission of forces, the socket must maintain an intimate fit with the

residual limb. This will provide a more natural feel as well as distribute the forces

throughout the limb (Itoga et al., 2013).

In the final design, both electromagnets and permanent magnets were spaced at

regular intervals around the outside of the socket. MAGNOLEV 1.0 consists of a

cylindrical plastic socket embedded with permanent magnets spaced between 1.5 inches

apart. These magnets are designed to supply an attractive force to the liner, which in turn

produces a negative pressure on the limb. Additionally, these magnets are capable of

resisting a small amount of shear force. This helps to keep the liner from rotating within

the socket when torsional forces were applied. These outer magnets would be used to

38

create the negative pressure on the sleeve. This utilizes the same mechanism as the

vacuum socket but relies on magnets to supply the outward force instead of a vacuum.

Permanent magnets would create the majority of the force required to generate the

desired negative pressure while electromagnets will be used to allow the user to adjust

how much negative pressure is applied. Inside of the socket a strong electromagnet is

placed at the center to create the holding force to attach the socket to the liner (Itoga et

al., 2013).

At the base of the socket, a circular hole was cut at the bottom so that a large

electromagnet could fit through it, as shown in Fig. 1.15 (a). When activated, the magnet

adheres to a metal plate on the bottom of the liner to create the suspension force. The

electromagnet was wired to a battery and switch, all of which was encased in a PVC pipe

located between the foot and the socket. The switch allowed for an on/off feature that

provided a release mechanism for the user (Itoga et al., 2013).

39

Fig 1.15 (a) Bottom view of socket (b) Side view (c) Front view

Fig. 1.15. Different views of prototype made by group in ME696

To attach the foot to the socket and hold the PVC housing in place, we passed screws

through a metal plate at one end of the pipe into holes at the bottom of the socket. Rather

than gluing or welding the pieces in place we chose screws to mount the PVC housing

because it allowed access to the battery pack in case the batteries required changing

(Itoga et al., 2013) Figures of the socket are shown in Fig. 1.15.

1.7 Thesis Organization

The ultimate goal of this research is to create an efficient technique for prosthetic

limbs that would compensate for residual limb shrinkage. Throughout the day, the

residual limb of an amputee can change in size. As a result, the prosthesis may become

uncomfortable to wear and may cause injuries to the residua limb.

40

1.7.1 Design and Prototype of the New Electromagnetic Suspension System

Using the MAGNOLEV as a start-off point, the first step was to modify the existing

prototype to create a more effective socket design. Because the focus of this project is on

the socket, the shaft and foot portions were disconnected. Itoga et al. (2013) had already

established that permanent were not the best of choice, therefore, electromagnets were

used as the main source of generating negative pressure. Several different designs of

electromagnet placement were created until one was officially chosen.

1.7.2 Development of the Feedback Control of Electromagnets by Pressure Sensor

Output

One of the major issues was the attractive strength of the electromagnets. In order to

test the highest amount of negative pressure the electromagnets could yield on the liner,

testing of the electromagnets were done with the permanent magnets and with a flat steel

bar. It was hypothesized that the attraction force of the permanent magnet with the

electromagnet would be stronger than steel bar and electromagnet.

The control system including the force sensors, microprocessor, motor controller, and

electromagnets had to be assembled and optimized. The challenges of the system are the

size, weight, power supply, and efficiency of the system.

1.7.3 Mechanical and Functional Testing

After a final prototype was created, testing had to be done to determine its

effectiveness. The experimental results were then compared with the parameters of an

existing vacuum suspension type prosthesis.

41

Chapter 2: CONCEPT AND DESIGN OF ELECTROMAGNETIC SUSPENSION

SYSTEM

2.1 Introduction

In this chapter, the design of the electromagnetic suspension system is described.

The design must be lightweight and comfortable while preventing volume limb loss. In

earlier work, a stack of eleven 5/8” diameter neodymium permanent magnets was shown

to generate a force of 3.19 N against an optimized liner consisting of seven layers of

magnetic paint and two layers of ferrous iron filings (Itoga et al., 2013). This proves that

magnets could be used as a feasible suspension technique. By spacing the stacks apart by

1.5 inches, a pressure of 2.196 kPa can be generated across the entire liner. Gershutz et

al. (2010) has shown that negative pressures between 2.0-3.3 kPa are sufficient to prevent

residual limb volume loss. This pressure of 2.196 kPa is well within the range necessary

to prevent residual limb volume loss. However, having stacks of magnets on 44 areas of

the socket can increase the size and weight of the prosthesis. Also, permanent magnets

are constantly “on”, making it not adjustable.

This electromagnetic suspension system builds on the principles proven by

vacuum systems to prevent limb loss and maintain a snug fit. Mechanically, an

electromagnet is made of a conductive wire, usually copper, that is wrapped around a

piece of metal several times. When a current is introduced from either a battery or other

source of electricity, it flows through the wire. This creates a magnetic field around the

coiled wire, magnetizing the metal. The strength of the magnet is directly related to the

number of times the wire coils around the rod. For a stronger magnetic field, the wire

should be wrapped more tightly. The tighter the wire is coiled around the core, the more

42

“loops” the current has to travel around it. In addition, the material used for the core can

also control the strength of the magnet. Electromagnets are useful because you can turn

the magnet on and off by completing or interrupting the circuit, respectively. (Brain and

Looper, 2000)

The magnetic attraction between the liner and the socket will create negative

pressure between the liner and the limb. This would make it easier for the wearer to

control the amount of force generated and, thus, the negative pressures exerted on the

limb. However, the proposed technology is smart, i.e. electromagnets and pressure

sensors will be strategically positioned throughout the socket where the limb is expected

to shrink with the ability to vary the attractive force between the liner and socket. The

result will be a suspension system that can vary the overall direction and magnitude of

the anchoring and suspension force to accommodate the user needs.

2.2 Socket Design

Keeping with the type of electromagnets chosen by the undergraduate group,

electromagnets were purchased from the company APW Company. Each electromagnet

had 1 inch diameter, a thickness of ¼ in and a holding force of 25 lbs, as shown in Fig.

2.1 (a) (Model EM100-6-222, APW Company). Each electromagnet was strategically

mounted at different points in the prosthetic where the limb volume loss would be

anticipated. To prevent pistoning, a ring of six electromagnets were placed at the top of

the socket. A ring of six electromagnets were also placed at the distal end of the socket in

order to prevent compression in this area. Four electromagnets were spread out in the

middle of the socket to supply further support around the limb. However, due to the

43

parameters of the battery packs, only three of the four middle electromagnets could be

connected to a battery source.

To mount the magnets on to the socket, a hole was drilled into the socket at the

site of each electromagnet. From the outside of the socket, the electromagnet was pushed

in just enough so that the edge of the electromagnet was flushed with the inside of the

socket. Each electromagnet was held in place with a custom bracket made from

galvanized steel. Once the electromagnet was positioned in a way that nothing could be

damaged (i.e. wires don’t get pinched), the bracket was mounted on with the

electromagnet using rivets and a bolt in the middle, securing it in place. This process is

shown in Fig. 2.1 (b)-(d). To organize the wires from the electromagnets, plastic zip-ties

were used to hold the wires together. This is shown in Fig. 2.1 (e). In order to simulate an

amputated limb, the bottom of the MAGNOLEV socket was cut out and replaced with a

PVC cap, as shown in Fig. 2.1 (f).

(a) (b) (c)

44

(d) (e) (f)

Fig. 2.1. Construction of the Prototype Socket

2.3 Liner Design

In order for the electromagnets on the socket to attract the liner, some kind of

magnetic material needed to be incorporated with the liner. By adding a magnetic

material to the liner, a secure attachment can be made when the electromagnets are turned

on. Testing of the electromagnet attraction forces were done with neodymium permanent

magnets and with a flat steel bar on a dynamometer machine.

A dynamometer is a device that can be used to measure force, torque, or power.

These compact units utilize the piezoelectric effect to acquire the highly dynamic signals.

There are basically two types: stationary and rotating dynamometers. With stationary

dynamometers, the object to be machined is mounted onto the stationary dynamometer.

These types of dynamometers are usually mounted on the machine table in order to

record cutting forces. Rotating dynamometers involve a rotating tool. To measure the

forces and torque in these applications, the rotating dynamometer is mounted directly on

the spindle of the tool. The dynamometer therefore rotates with the tool. Stationary

piezoelectric dynamometers output a charge linearly proportional to the acting load. The

45

charge amplifier converts this charge into an analog voltage signal, which is then

recorded by the inline data acquisition unit (Kistler, 2014).

The dynamometer used was the 4-Component Dynamometer for Cutting Force

Measurement in Drilling - Type 9272 – by Kistler. It is a static piezo-multicomponent

dynamometer consisting of just one four-component sensor, which is mounted and

preloaded between a base plate and a top plate. The Kistler 9272 is able to measure four

components – forces in the x-, y-, and z-direction as well as the torque in the z-direction.

In each direction (x, y, and z), the Kistler 9272 had a measuring range of -5 to 5 kN.

However, when measuring the torque, the measuring range was only -0.2 to 0.2 kN·m.

The dynamometer is 70 mm in height, has an outer diameter of 100.0 mm, and an inside

diameter of 15.0 mm (Kistler, 2014).

Fig. 2.2. Kistler 9272 Dynamometer (kistler.com)

Data from the dynamometer was set to record every 10 MU/Volt, 3.50 PC/MU

and the output was graphed through a LabView program on a connected computer. It was

hypothesized that the attraction force of the permanent magnet with the electromagnet

would be stronger than steel bar and electromagnet. The concept of the test was to have

an electromagnet held in place with the dynamometer (at a constant voltage) and to

46

determine which required more force to pull away from the electromagnet – the

permanent magnet of the steel bar. Several trials were done for each and the absolute

value of the average force was calculated. Besides the first trial, it was shown that the

attraction force with the steel bar was stronger. As a result, pieces of the steel bar were

incorporated when creating the new liner design. Results of this test are shown below.

Table 4: Force Required to Pull Apart – Electromagnet/Permanent Magnet02

(Settings: 3.50 PC/MU, 10 MU/Volt)

Trial No. Time

(s)

Dynamometer Output (V) Force to Pull

Electromagnet/Permanent Magnet

Apart (N)

1 4.305 -1.06656 -10.6656

2 7.222 -1.07661 -10.7661

3 10.426 -1.08859 -10.8859

4 13.002 -1.14335 -11.4335

5 15.624 -1.10577 -11.0577

6 18.619 -1.17024 -11.7024

7 21.463 -1.23893 -12.3893

8 24.335 -1.21916 -12.1916

9 27.43 -1.28623 -12.8623

10 30.398 -1.13881 -11.3881

11 33.125 -1.33385 -13.3385

Average = 11.698 N

47

Table 5: Force Required to Pull Apart – Electromagnet/Steel Bar02

(Settings: 3.50 PC/MU, 10 MU/Volt)

Trial No. Time

(s)

Dynamometer Output (V) Force to Pull Electromagnet/Steel

Bar Apart (N)

1 6.561 -1.52274 -15.2274

2 10.012 -1.77902 -17.7902

3 12.785 -1.36593 -13.6593

4 17.254 -1.71325 -17.1325

5 21.01 -1.76638 -17.6638

6 24.628 -1.89986 -18.9986

7 28.49 -1.97795 -19.7795

8 31.345 -1.68571 -16.8571

Average = 11.714 N

Initially, the steel pieces were placed inside of the liner. However, due to the

amount of space between the electromagnet and liner, there was no attraction force and it

was determined that the steel piece could not be installed inside the liner. To reduce the

amount of space between, it was established that the steel would stay on the outside of

the liner. In order to create the strongest points of attraction, the steel pieces were glued

onto the liner in areas where it would be in direct contact with the electromagnet on the

socket (areas were marked before material was glued on).

48

Fig. 2.3. Construction of Prototype Liner

2.4 Power Supply

In order to power the electromagnets and the electrical components, a power

supply was needed. Each electromagnet required a voltage of 6V to have maximum

attraction force. Two options were considered – using one large Duracell 6V lantern

battery or by combining 4 1.5V AA batteries together. Due to space and weight

restrictions, using 4 1.5V AA batteries together was the better option. Through research,

it was found that the maximum amperage one battery pack (holding 4 AA batteries) may

hold is 2 amps. Since each electromagnet emits 0.66 amps, 3 electromagnets could be

connected to one battery pack. As a result, 2 battery packs were needed for the 6

electromagnets at the top ring on the socket. In the middle of the socket, 4

49

electromagnets were placed. Due to the parameters of the battery pack, only 3

electromagnets were connected to a power source, leaving the last inactive.

50

Chapter 3: MECHANICAL PERFORMANCE

3.1 Introduction

After a socket and liner design were finalized, testing was done to determine the

attractive force between the socket and liner with the electromagnets activated. An

Instron testing machine was used in tension to measure the force between the socket and

liner connected by an array of electromagnets. Though testing was done in tension, the

purpose was to establish if there was a force capable of breaking the attachment between

the steel pieces on the liner and the electromagnets.

3.2 Instron Testing Set-Up

Testing of the socket and liner was done using an Instron machine – Model 4206.

Instron universal testing instruments are highly reliable precision systems for evaluating

the mechanical properties of materials. The advantages to these systems, as accurate and

versatile tools, make them equally adaptable to research and development requirements

as to repetitive testing applications of production quality control (Model 4206 Operators

Guide).

The Model 4206 is an electromechanical device based on strain gauge load cells

and servo-control systems. It is comprised of two major assemblies. The first assembly

consists of a crosshead driver and control system, which applies tensile or compressive

loading to specimens. The structure of the Model 4206 consists of two vertical

leadscrews, a moving crosshead, and a baseplate which is the lower bearing carrier for

the leadscrews. Additional structure includes two columns that enclose the leadscrews

and a fixed top plate which is the leadscrew upper bearing carrier. The entire assembly is

installed on a base which encloses the mechanical drive train for the leadscrews, a servo

51

amplifier, and a connector panel which provides interfacing with the control console

(Model 4206 Operators Guide).

The mechanical drive train for the moving crosshead consists of a DC servomotor

coupled by a timing belt through a belt box containing a series of timing belts, pulleys,

and clutches to the leadscrew drive pulleys. The control of the moving crosshead is

developed from a commanded testing speed and direction, programmed at the console.

The encoder is an optical device that provides feedback information by pulses which

indicate displacement of the crosshead. This signal is applied to the crosshead control

board in the console (Model 4206 Operators Guide).

The second assembly consists of a highly sensitive load weighing system which

measures the loading of a specimen. Instron load cells are precision force transducers

containing strain gauges bonded to internal load bearing structures which are stressed

during a material test by applied tension or compression forces. Three basic design

structures are used for the load cell which include bending beam, axial stress, and shear.

The load cells function with the highly sensitive load weighing system incorporated in

Instron testing instruments. The element in each load cell includes a balanced bridge

employing strain gauges supplied with excitation from the signal conditioner. When a

load is applied to the cell, the element is deformed and the bridge becomes unbalanced.

The result is the generation of an electrical signal in direct proportion to the deformation

of the element. The highest position the Model 4206 is able to reach is about 52 inches.

At the top of the machine is a mounted load cell with a maximum loading capacity of 150

kN (Model 4206 Operators Guide).

52

The basic operation of the instrument consists of selecting a load cell for a

particular application, mounting the cell in the moving crosshead within the load frame,

then setting a specimen in position so that an applied load can be measured. Two grips at

the top and bottom hold the specimen in place. For tensile testing, the fixtures move away

from each other, stretching the specimen. Tensile, or compressive if applicable, forces are

applied when the moving crosshead is operated by two vertical leadscrews (Model 4206

Operators Guide).

Before testing was started, modifications on the socket needed to be done. At the

bottom of the socket, a hole was drilled through so that a small “peg” could be screwed

in. This peg would be used to hold the socket in place inside the clamps of the Instron

machine (Fig. 3.1). To simulate a leg being in the socket, a round fixed head was used

and placed inside the liner, as shown in Fig. 3.2. Ropes were used to secure the top part in

place inside the top Instron clamps. Instead of folding the liner down, it was at its full

length. This method proved to be successful. Testing was done to establish if there was a

force capable of breaking the attachment between the steel pieces on the liner and the

electromagnets.

Fig. 3.1. Attachment of “peg” to Socket

53

A total of 9 electromagnets at the top and middle were powered at a constant 6V.

The speed of the machine was set to move at 3 mm/min Using the Bluehill 2 software,

the amount of extension (mm) and the amount of force generated (N) were recorded.

Fig. 3.2. Instron Testing Set-Up

3.3 Results and Discussion

Fig. 3.3. Time vs Load Plot from Instron Testing

54

According to Gerschutz et al. (2010), a pressure of 10-15 mmHg is sufficient

enough to prevent shrinkage of the residual limb and may even promote limb health. In

this single subject study, Gerschutz et al. focused on a K2 transtibial amputee new to

elevated vacuum suspension with a history of residual limb volume fluctuations. They

investigated the effects on the limb at different vacuum pressure settings at three

treatment levels: absence of elevated vacuum (suction), vacuum (negative pressure) at 10

in Hg, and vacuum (negative pressure) at 15 in Hg using the LimbLogic® VS

technology. The results indicated a significantly less volume fluctuation with vacuum

compared with suction and an improvement in volume retention with continued vacuum

usage.

10-15 mmHg of pressure can be converted to 1.33-1.9998 kPa of pressure. Using

the formula P=F/A (where P is pressure, F is force, and A is area), the amount of force

that needed to be generated from the electromagnets to compensate the required pressure

could be found. The liner given by Mr. Newton had a circumference of 30 cm, which was

measured about 5 cm from the distal end. Knowing this information, the radius was found

to be 0.0477 m and the area was calculated out to be 0.014 m2. By knowing the pressure

and area, the force needed to generate 1.33-2.00 kPa of pressure was found to be 19.01-

28.589 N.

After reviewing the data, it was shown that a force of 19.01-28.589 N could be

generated by the electromagnets. During testing using the third set-up, the highest force

that could be generated was 29.01N. However, during testing, the electromagnets were

strong enough to hold the liner in place and it was seen that the liner did not move. As a

result, it could be concluded that more force could be generated until the socket and liner

55

reached its yielding point. It should be noted that there was negligible force generated

from friction between the liner and socket because the setup did not include anything that

would create a normal force between the liner and socket similar to the leg of the user.

3.4 Conclusions

The Instron testing was done to determine if the top and middle electromagnets

generated enough attraction force to hold the liner in place with the socket. After review

the results, it is clear that the design is adequate enough to reduce the issue of pistoning

and compression around the middle areas of the limb. Since the yield value was not

reached during the test, it is possible that the design would be able to withstand even

more force.

56

Chapter 4: FUNCTIONAL TESTING

4.1 Introduction

The foundation of this new suspension system is the array of pressure sensors,

electromagnets, and magnetically attractive material positioned throughout the socket and

liner. In this chapter the system that controls the electromagnets by pressure sensor output

is described. The system, which includes a feedback control loop, was assembled and

tested to ensure proper function. The results of this testing are also presented in this

chapter.

Due to the fact that each limb has a different composition, it is difficult to

determine which part of the limb will need more attractive force. However, excessive

amounts of compressive force at the distal end of the limb results in limb damage. As a

result, the bottom ring of electromagnets were designed to have adjustable attraction

forces. These force sensors will individually control a bottom ring of six electromagnets.

4.2 Control of Electromagnets

Figure 4.1 shows the diagram for the closed loop control system. The system

output is the electromagnet force. The force sensor monitors the system output (the force

between the electromagnet and liner) and feeds the data to an Arduino controller which

adjusts the electrical signal to the motor driver as necessary to maintain the desired

system output. When the limb shrinks and the liner pulls away from the socket, the

decrease in force is measured, and the signal changed to increase electromagnetic force,

pulling the liner back to the socket. Feedback from measuring the force allows the

controller to dynamically compensate for changes to the force.

57

In the circuit the force sensors were connected to an amplifying circuit and an

Arduino microprocessor was used to read the amplified signal from the pressure sensor.

The Arduino microcontroller would adjust the attraction force of the electromagnets,

therefore varying the forces exerted on the limb. Since the sensor output was a small

amount of voltage, an amplifying circuit was also incorporated into the circuit. By

supplying a constant voltage of 4.5V to the force sensors and amplifier a reasonable

output was created. At the moment, two external power sources were used. One was used

to power the force sensors while the other was used to power the amplifier. A diagram of

the connected electrical components is shown below. The force sensor generates a small

charge that must be amplified by an inverting amplifying circuit. The amplified signal

was fed to an Arduino microprocessor (Model ATmega2560) which controlled the motor

driver to power the electromagnets.

Fig. 4.1. Feedback Loop of Electrical Components

4.2.1 Tekscan Force Sensors

A pack of 8 Standard FlexiForce Sensors – Model A201 were purchased from the

company Tekscan. FlexiForce sensors are ultra-thin and flexible printed which can be

easily integrated into force measurement applications. FlexiForce sensors are

piezoresistive sensors that can be utilized in many ways, are durable, and can be made

58

into a variety of shapes and sizes. The FlexiForce sensor acts as a force sensing resistor in

an electrical circuit. The maximum voltage applied to the sensors should not exceed 5V.

The maximum power required to run the force sensors is 0.0125 watts (5V, 2.5mA).

When the force sensor is unloaded, the resistance is very high. When a force is applied to

the sensor, the resistance decreases. Other benefits to the FlexiForce sensors are that they

have superior linearity and accuracy and are able to handle a large range of forces and

temperatures (FlexiForce Sensors).

The standard A201 FlexiForce sensors are a 3-pin male square pin sensor. They

are constructed of two layers of substrate (polyester) film. The active sensing area is

defined by the silver circle, which is designed to be on top of the pressure-sensitive ink.

The male square pins at the end of the sensor allow the sensor to be easily incorporated

into a circuit. Model A201 FlexiForce sensors have a thickness of 0.203mm, a length of

191mm, and a width of 14mm. The sensing area is 9.53mm in diameter. These sensors

can sense forces ranging from 0-100 lbs (FlexiForce Sensors).

Fig. 4.2. Tekscan FlexiForce Sensor -Model A201 (Tekscan.com)

59

Fig. 4.3. Components of Tekscan FlexiForce Sensor (Tekscan.com)

To create the circuit of these sensors, several components were needed. At bottom

pin, a wire was used as the connection between the circuit and the power source. Since

the force sensor required power to function, an external power source was used. The

center pin was inactive and was therefore left alone. At the top pin, several more wires

were required. Also, in order to amplify the outputted signal, an operational amplifier

(MCP 6004) was connected. The wires at the top pin were used as a way to ground the

circuit and as a connection for a power source for the amplifier component. A resistor

was also connected and used as a reference resistor. This resistor was used to make sure

that the circuit would remain stable and accurate after calibration.

These sensors were chosen because of their long thin design. Since they are thin

and flexible, they can be easily incorporated inside of the socket, minimizing the distance

between the force sensor and limb, and can conform to the shape of the limb. Also,

60

because they are long, they can be placed at the top of the socket but still be connected to

the circuit located at the bottom.

4.2.2. Arduino Microprocessor

In order to make the electromagnets adjustable, the circuit was connected to a

microcontroller, Arduino (model ATmega2560) as shown in Fig. 4.4. The Arduino Mega

2560 has 54 digital input/output pins (15 of which can be used as pulse width modulation

(PWM) outputs), 16 analog inputs, 4 hardware serial ports, a 16 MHz crystal oscillator, a

USB connection, a power jack, an ICSP header, and a reset button. Arduino was powered

via a USB connection and required a minimum of 5 volts of power to function correctly

(Arduino Mega 2560). The pressure sensors were mounted onto a circuit board with an

amplifier connected to the Arduino microprocessor. The microprocessor read the output

from the pressure sensor to drive the electromagnet.

Pulse width modulation (PWM) is a method for generating an analog signal using

a digital source. The width of an impulse conforms to the signal information. PWM is

used to control the amount of power being supplied to the electrical system. A PWM

signal consists of two main components that define its behavior – a duty cycle and a

frequency. The duty cycle describes the amount of time the signal is in a high (on) state

as a percentage of the total time it takes to complete one cycle. The frequency determines

how fast the PWM completes a cycle, and therefore how fast it switches between high

and low states (PWM). For the electromagnetic suspension system, two cases will happen

depending on the PWM signal from Arduino. The first case will be that if a low amount

of force is detected by the sensors, this signifies that there is not a secure attachment at

the liner and socket interface. As a result, the attractive force of the electromagnet will

61

increase to pull the liner closer to the socket wall. On the other hand, the second case will

be that if a high amount of force is detected, this signifies that there is a large amount of

force compressing against the socket. As a result, the attractive force from the

electromagnet will decrease to ease the amount of pressure exerted on the limb.

Therefore, as the amount of force increases, the width of the duty cycle will decrease.

With help from the Arduino program, AnalogInOutSignal, these magnets could be

adjusted. The program AnalogInOutSignal read the signal from an analog input pin,

mapped the results to a range from 0 to 255, and used the result to set the PWM of an

output pin. It also printed the results to the serial monitor for verification.

As stated above, when force was applied to the sensor, the Arduino reads the

results and inversely modify the attraction force of the electromagnet depending on how

high the force is (as force increased, electromagnet attractive force decreased). However,

because the output reading gives out a value that is double the input, the Arduino code

needed to be calibrated to the force sensor. The modified Arduino code is shown in Fig.

4.8.

Fig. 4.4. Arduino Model ATmega2560 (Arduino.com)

62

Fig. 4.5. Circuitry of Force Sensors and Arduino

4.2.3 Rover 5 Motor Driver

In the case involving the amplifier, it was noticed that the amount of voltage

exerted after Arduino had changed was still too low. As a result, a motor driver was

connected to the Arduino board. The motor driver selected was the Rover 5 Motor Driver

Board by Sparkfun. Originally designed for any small 4-wheel drive robotic vehicle, the

Rover 5 motor driver comes with four motor outputs, four encoder inputs and current

sensing for each motor. The motor drivers can be controlled by simply applying a logic 0

or 1 to the direction pin for that motor and a PWM signal to the speed pin (both which

could be connected from the Arduino board). (Rover 5 Motor Driver Board)

There are two power connectors on board - one is for 5V logic and the other is the

motor supply. The board is rated for a maximum motor supply voltage of 12V. As a

result, two more battery packs needed to be connected for the motor driver. The built in

63

control logic allows each motor to be controlled by 2 pins. Driving the direction pin high

or low will cause the motor to run forward or reverse. The PWM pin is used to control

the motor speed. When this pin is low, the motor is off. When this pin is high the motor is

at full power. To vary the speed of the motor this pin must be Pulse Width Modulated

(Rover 5 Motor Driver Board). Depending on the output voltage from the pulse width

modulation of the Arduino (which is dependent on how much force is exerted on the

sensor), the amount of attractive force exerted by the electromagnet was increased

substantially.

Fig. 4.6. Rover 5 Motor Driver Board (Sparkfun.com)

64

Fig. 4.7. Circuitry of Force Sensor, Arduino, and the Motor Driver

After establishing that these connections do work, the focus was to change the

Arduino code to match the concept of this project – when the pressure decreases, the

microprocessor increases the power to the electromagnet to pull the liner, and thus the

tissue, toward the socket. If the sensor detects a low range of force, it means that the limb

is not properly in place in the socket and that the electromagnets should increase their

attraction force to the liner. This could easily be done by modifying the Arduino code as

shown below.

// These constants won't change. They're used to give names

// to the pins used:

const int analogInPin = A0; // Analog input pin that the potentiometer is

attached to

int sensorValue = 0; // value read from the pot

int outputValue = 0; // value output to the PWM (analog out)

void setup() {

// initialize serial communications at 9600 bps:

Serial.begin(9600);

pinMode(9,OUTPUT);

}

void loop() {

// read the analog in value:

65

sensorValue = analogRead(analogInPin);

// map it to the range of the analog out:

outputValue = map(sensorValue, 0, 500, 0, 100);

// change the analog out value:

digitalWrite(9, HIGH);

delay(100-outputValue);

digitalWrite(9, LOW);

delay(outputValue);

// print the results to the serial monitor:

Serial.print("sensor = " );

Serial.print(sensorValue);

Serial.print("\t output = ");

Serial.println(outputValue);

// wait 10 milliseconds before the next loop

// for the analog-to-digital converter to settle

// after the last reading:

}

Fig. 4.8. Arduino Code

4.3 Experimental Setup for Testing the Circuit Operation

An experiment was setup to test the circuit operation. In all, testing required two

computers, two external power sources, the force sensor circuit, the Arduino

microprocessor, the motor driver, and two battery sources for the motor driver. A

panoramic view of these components is shown in Fig. 4.9.

Two computers were used for testing. One of the computers (the laptop) was

connected to the Arduino microprocessor. Not only did this provide power to the

microprocessor, but if any modifications to the code needed to be made, it could be done

using the laptop. The second computer (the desktop) contained the LabView program

needed to output the results.

Between these two computers, the Arduino board and motor driver were

connected together. Two external power sources were used – one to power the force

sensors and one to power the amplifier in the circuit. In order to power the motor driver,

two external battery packs were connected.

66

For the experimental procedure, the end of the force sensor was placed onto the

dynamometer. The end of a pencil was used to apply force to the sensing area because the

shape was a good fit for the force sensor and area could be easily calculated. The force

was measured by the dynamometer and the voltage signal outputs from the Arduino

controller and motor driver were measured as indirect measures of the electromagnetic

force. The measurements were recorded by National Instruments data acquisition

hardware and LabView software.

Fig. 4.9. Panoramic View of Circuitry Testing Set-Up

Motor Driver and Electromagnet

Computer w/ LabView

Computer to modify Arduino

code

Dynamometer with pencil pressing on sensor

External power supplies

FlexiForce sensor circuit

Arduino

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4.4 Results and Discussion

In order to establish that the Arduino code and motor drivers were working

correctly, testing was done. This would determine how much force the sensors could

handle to reach a certain voltage. This certain voltage would cause the varying attractive

force of the electromagnets. Once all the components were connected, testing was done

using the dynamometer to quantify the correlation between force sensor output and

electromagnetic attractive force.

4.4.1 Arduino Output

When pressing down on the sensor, an object with a constant area was needed.

Therefore, a pencil with a diameter of 0.735cm was used. Still, due to the sensitivity of

Arduino, when the sensor value exceeded the set maximum value of 100, the Arduino

and motor driver board malfunctioned.

Due to the parameters of the dynamometer, the forces pushing down onto the

sensor must be multiplied by 10N. The highest force recorded by the dynamometer (after

multiplying by 10) was 30.98 N. This result meant that 31 N of force pushing on the

force sensor is required for the electromagnet to output the least amount of attractive

force it can. As seen from the results, when the amount of force is low, the electromagnet

is on (voltage at 6V). As more force pushes down on the sensor, the Arduino causes the

width of the PWM to widen. When the maximum amount of force is pushed down onto

the sensor, this results in the voltage being mostly at 0V (mostly off).

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Fig. 4.10. Arduino Output Results

Fig. 4.11. Arduino Output: Close-Up View at Low Force

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Fig. 4.12. Arduino Output: Close-Up View at an Increasing Force

4.4.2 Motor Driver Output

The motor driver was attached in order to amplify the signal from the Arduino.

With the motor driver board, the attractive force of the electromagnet increased. As with

the Arduino, when there is no force, the PWM from Arduino is mostly on at about 6V

and the electromagnet has maximum attractive force. However, as the force on the sensor

increases, the width between each cycle in the PWM widens, establishing that the

electromagnet is losing its attractive force as more force is being applied onto the sensor.

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Fig. 4.13. Motor Driver Output Results

Fig. 4.14. Motor Driver Output: Close-Up View at Low Force

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Fig. 4.15. Motor Driver Output: Close-Up View at Increasing Force

4.5 Conclusions

When there is a high amount of force pressing against the limb in the socket, the

limb may become injured and the socket fit will feel uncomfortable. With the Arduino

and motor driver board, the electromagnets on the socket will be able to adjust

automatically if too much force is being applied or not. By looking at the results, this

concept of a small electromagnetic suspension system consisting of force sensors, a

microprocessor, motor driver, and electromagnets is possible to create and provides

adequate force for attachment of the liner to the socket.

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Chapter 5: CONCLUSIONS AND FUTURE WORK

5.1 Summary

Figures 5.1 and 5.2 show the initial setup of the components that make up the

prototype: electromagnets mounted to the socket, battery packs, Arduino microprocessor,

and pressure sensors. Though not included in figures 5.1 and 5.2, there will also be a

motor driver board connected to Arduino. The pressure sensors will be placed through a

hole in the bottom of the socket and secured on top of the electromagnets to be

controlled.

Fig. 5.1. Initial Set-Up of Prototype Components (Sensors Inserted)

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Fig. 5.2. Initial Set-Up of Prototype Components (Sensors Not Inserted)

5.2 Conclusions

Though this research, it was shown that the array of electromagnets provided

adequate force for attachment at the liner and socket interface. A small electromagnetic

suspension system was possible, consisting of force sensors, a microprocessor, motor

driver, and electromagnets. It was also shown that this electromagnetic suspension

system was capable of varying attractive forces at different points based on the output of

the force sensors.

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5.3 Contributions

This research sets a completely new direction in the area of prosthetics and will

advance the knowledge and understanding of human-device interaction. With further

development, this research will have immediate social, economic, and scientific impact.

It will improve daily lives and enable a broader range of activity for amputees. This

research will create a much needed intelligent method for the suspension of prosthetic

limbs. Three specific contributions were made in this research:

1. The initial design and prototype of a prosthetic incorporating an electromagnetic

suspension system

2. Theoretical and experimental determination of mechanical performance of the

electromagnetic suspension

3. Development of a closed loop feedback control system for electromagnetic

suspension

5.4 Future Work

The relationship and control strategy between individual, local pressure

measurements and attractive force of controlled electromagnets must be completely

characterized and established. The control system, including power supply, must be

developed and efficiently integrated into the prosthetic. In order to accomplish this, the

sensitivity and accuracy of the sensors will need to be measured and characterized for

proper control of the electromagnets. A single pressure sensor-electromagnet feedback

loop will be constructed and tested.

After the system is established, the task will be to create a prototype of the new

prosthetic. The components must be integrated into the carbon fiber socket, silicone liner,

75

and limb of the prosthesis. The resulting prototype will be functionally tested and

optimized for phase 2 of the project. It should be noted that the type of pressure sensor

and electromagnets might change. The preliminary results have shown acceptable

variability of the pressure sensors, but a pressure sensor with better resolution and smaller

range might be more suitable for this application. Also, the electromagnets have

sufficient holding force, but a smaller electromagnet with less force might be adequate.

The weight, power, and thermal requirements of the control system must be

minimized. There is much work to be done to get the system to acceptable power, size,

and weight levels. As mentioned before, the electromagnets could possibly be smaller

and weaker for this application. The number and strength of the electromagnets will be

experimented to maintain proper balance between weight and anchoring strength of the

prosthetic. Also, heat builds when the electromagnets are used at full power for extended

periods of time, so the configuration and control system must be designed for intermittent

use of the electromagnets. The system is powered by AA batteries currently, but this will

probably be changed to lighter, more robust power supply. The final design must find

balance between the size, weight, and power requirements and minimum values for

anchoring strength and limb volume control.

Once the components have been finalized, they will be packaged into a case.

Then, the second prototype will be tested for strength, adhesion of the liner to the socket,

and power requirements. The prototype will be mounted in the Instron testing unit just as

before, but in different configurations. The system should use minimal power while

resting and keep the liner snugly fit in the socket under movement and force. The

pressure sensor and electromagnet outputs will be monitored during testing to gage the

76

function. Lastly, the parts need to be assembled into a working prototype that can be

worn and tested by an amputee.

77

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