Bio Mechanical

8/8/2019 Bio Mechanical

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Biomechanical Energy Harvesting 1

College Of Engineering ,Chengannur

1.INTRODUCTIONHuman power is an attractive energy source. Muscle converts food into positive mechanical

work with peak efficiencies of approximately 25%, comparable to that of internal

combustion engines. The work can be performed at a high rate, with 100 W mechanical

easily sustainable by the average person and twice that sustainable by elite athletes. Food,

the original source of the metabolic energy required by muscles, is nearly as rich an energy

source as gasoline and approximately 100 fold greater than batteries of the same weight.

Given these attractive properties, it is not surprising that a number of inventions have

focused on converting human mechanical power into electrical power. These include hand

crank and bicycle generators as well as windup flashlights, radios, and cell phone chargers.

One major drawback of these devices is that they require dedicated power generation by the

user. This serves to limit the time available to produce power and, thus, the amount of

useful energy that can be generated.

Biomechanical energy harvesters generate electricity from people as they go about their

activities of daily living. This results in power generation over much longer durations. An

exemplary energy harvesting device is the self-winding watch which produces enough

electricity to power the device without requiring the user to wind it but is insufficient for

most of our portable power needs. There are a number of devices based on the same

fundamental principle as the self-winding watch using an external load to drive a generator.

The most successful design to date is the spring-loaded energy harvesting backpack that

converts the packs linear motion relative to the user into rotational motion of a rotary-

magnetic generator producing as much as 7 W .A second group of energy harvesters use the

bodys own inertia to generate electricity from the compression of the shoe sole harvesting

as much as 0.8 W. Our technique differed from other techniques in two main ways. First,

the device took advantage of the fact that much of the displacement during walking occurs

at body joints and harvested energy from knee motion rather than from an external load or

the compression of the shoe sole. Second, the device selectively engaged power generation

to assist the body in performing negative work, analogous to regenerative braking in hybrid

cars. The main purpose of this seminar is to explain the physiological principles that guided

our design process. We also present a brief description of our device design and its

performance.


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2.DETAILED DESCRIPTION2.1 THE COMPRESSION OF THE SHOE SOLE

2.1.1 Overview of Design

The predominantly compressive forces of a heel strike have been harnessed to provide the

tensile forces best suited to excite the PVDF by means of a heel insert. The advantage of

situating the transducer beneath the heel of the foot instead of farther forward near the ball,

lies in the fact that there is more energy dissipated in this location. The wearer's body

weight initially falls wholly on the heel and is only gradually transferred forward with the

step. The heel insert was constructed around a horseshoe-shaped piece of rubber material

cut out from the heel of a sneaker. Two horizontal heel-shaped polycarbonate plates were

glued at their curved edges to the top and bottom rubber edges of the shoe's heel cutout.

Fifteen elongated, rectangular unimorph strips were in turn glued vertically between the two

plates, along shallow front-to-back grooves cut in the polycarbonate. The rubber cutout

erves both to protect the strips from excess compression, i.e. to dissipate the forces which

the strips do not absorb, and to maintain the natural feel of the shoe. Indeed, there is little to

no difference in the sensation of walking with the transducer mounted in the sneaker.

Fig(2.1).Shoe insert composed of(1) Rubber cutout,(2) Poly carbonate plates,(3) copper terminals and (4)

unimorph strips

The unimorph strips themselves were each constructed of one 0.5 inch tall, 52um thick

silver laminated PVDF film (Measurement Specialties) bonded with cyanoacrylate to the


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side of a slightly wider and longer 4 mil thick PET plastic film substrate (the strips vary in

length with the changing space available in the cutout from 1 to 2.25 inches). This particular

substrate was chosen for its stiffness and spring-like qualities following much

experimentation with different materials. Other plastics failed to return to their original

shape after deformation, began to craze, or were too thick to be bent under a reasonable

force. During a heel strike, the polycarbonate plates are compressed together, in turn

bending all of the PET strips aligned between them. The bending plastic strips induce a

strain in the bonded PVDF film, which is offset from the neutral axis. Compared to bending

solitary strips, this unimorph configuration substantially increases the electrical response of

the PVDF.

Each piece of film was glued to the substrate in the same orientation, with the stretch

direction aligned vertically and the positively poled side of the film facing away from the

substrate. A narrow copper wire was bonded to the inner electrode of each laminate with

conductive epoxy. Another wire was taped to the outer electrode. These leads from each

strip were connected to two copper terminals on the outer edge of the shoe insert so that all

of the charge generators would act in a parallel configuration to maximize the generated

current.

The back ends of the unimorph strips were purposefully cut slightly too large to fit perfectly

straight between the two polycarbonate plates, causing them to remain very slightly bent

when the shoe insert is not in compression. This pre-bending ensures that each of the strips

bends in the same direction under compression so that each piece of film undergoes a tensile

strain and the sign of the voltage produced from each strip is equivalent. This is important

because a strip bending out of unison will cancel the voltage produced by another,

decreasing the effectiveness of the transducer.


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2.1.2 Power And Efficiency

Two methods were utilized to determine the power output of the transducer system. In the

first case, the potential drop across a known load resistor was measured. In order to

determine the resistor best matched to the impedance of the transducer, the power output for

various loads was measured while the transducer was compressed under a constant force.

The results from this experiment, depicted in Fig(2.2), show that the power peaks at load of

about 500 k .

Fig (2.2) Dependence of power on load resistance, showing best resistance match

The fairly high resistance required to match the transducer's impedance can be explained by

imagining an equivalent circuit composed of the piezoelectric charge generator, a capacitor

for the system's internal capacitance, and a resistor to model the dielectric leakage across

the PVDF. For low frequency applications, however, the internal film resistance is veryhigh and can be ignored. The structure's low net capacitance of 17 n F requires that the

matching resistance be large. The best load value determined, the voltage from the

transducer across a 470 k_ resistor at a 1 Hz frequency, shown in Figure 3, was recorded

and the root mean square voltage calculated, giving an average power of 0.06 m W. The

peak voltage was measured at 21 V, for a peak power of 0.94 m W. The sharp initial peaks

are caused by the fairly coherent compression of the strips as the weight of the body falls on

the heel.


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The subsequent rolling transfer of weight forward towards the ball of the foot allows the

strips to straighten, generating the ensuing oppositely signed voltage. The negative spikes

between steps can be attributed to individual strips returning to their fully extended

positions and also to small tensile forces within the heel.

Fig (2.3) Voltage wave form transducer with 470 k load

In the second method, the energy stored on a bucket capacitor was calculated. The capacitor

was connected to the transducer circuit through a full-wave bridge rectifier. With a 1 F

capacitor (1 Hz excitation), the peak voltage reaches 9.6 V over a 1 second period, yielding

an average power of 0.05 mW. This result is slightly lower than that found with the resistive

load because the 1 F capacitor does not as closely match the impedance of the transducer.Tests were also conducted with a larger 100 F capacitor to demonstrate the effects of a

larger storage medium. Predictably, the stored energy was about an order of magnitude

lower. This highlights the necessity for efficient power conditioning when converting the

output energy to a useful form.

In addition, the electromechanical efficiency of the entire shoe insert was calculated by

comparing the energy required to compress the strips inside the shoe insert with the energy

generated during a corresponding period. The net energy required to compress a single 1

inch long unimorph strip a distance of 1mm was 0.2 mJ. This value can be extrapolated to

the case of the complete transducer, for a total mechanical input of 5.9 mJ. Using the

average power generated by the transducer over a 1 second period, the efficiency of the

transducer is found to be approximately 1%. This measurement takes into account losses

caused by the imperfect compression of the multiple strips along with losses in the transfer

of power from the PVDF to the load. The PVDF stave developed by the MIT Media Lab

achieved an electromechanical efficiency of 0.5% [4]. This calculation, however, used the


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open circuit voltage from the transducer to determine the raw electrical output. This

approach proved to be extremely difficult in the case of the heel-mounted transducer

because its much lower capacitance was not large enough to sustain a voltage long enough

to be measured accurately. A dual unity gain buffer amplifier circuit was constructed of two

high power op-amps in an attempt to measure the potential but the maximum supply voltage

was not high enough to avoid capping. Very high power op-amps were not considered due

to financial considerations. An accurate measurement of the open circuit voltage would

theoretically give an approximate raw electromechanical efficiency of 2%.

2.2 THE SPRING-LOADED BACKPACK

2.2.1 Overview

Hikers toiling under the weight of a heavy backpack needn't just get hot and sweaty from

their efforts. Some of the energy they expend in walking can now be captured by a

backpack devised by US researchers, which converts it to electricity that can power portable

electronics.

Walking is a particularly good source of human power: during steady hiking the muscles produce up to 100 W. As far back as 1967, scientists at the Massachusetts Institute of

Technology used piezoelectric devices, which generate electricity when squeezed, inserted

into the heel of a shoe to create power for portable electronics such as pacemakers. But

mechanical generators housed in shoe heels have tended to be rather cumbersome and

fragile.

Fig (2.4).Spring loaded back pack structure


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The backpack is spring-loaded, so it can bounce up and down when the wearer walks. This

moves a toothed rod, which meshes with a gearwheel. As the gearwheel rotates, it generates

electricity.

The heavier the load, the more power the pack generates. But the researchers were

concerned that the shifting load might force the wearer to use more energy in walking,

much as a bicycle dynamo puts a greater strain on a cyclist. So they measured the metabolic

rates of their test subjects by looking at how much oxygen they consume.

2.2.2 Power And Efficiency

The pack, which weighs about 20-38 kg depending how much power you need, generated

up to 7.4 watts of power when tested on a treadmill. That's enough to keep your GPS locator

and a head-lamp running indefinitely in the wilderness - potentially useful for soldiers or

rescue workers. They could even take a break now and then without losing power, as

surpluss energy is stored in lightweight batteries.

2.3 KNEE DEVICE FOR ENERGY HARVESTING

When comparing the above discussed energy harvesting devices, it has many disadvantages.The Knee device overcomes the problem found in above two cases. The following section

gives the detailed design and power generation method of knee devices

2.3.1 Walking Biomechanics

To effectively harvest energy from walking, it is necessary to first understand walking

mechanics and the underlying muscle function. During walking at a constant speed on level

ground, no net mechanical work is performed on the body since there is no net change in

kinetic energy (i.e. speed) or potential energy (i.e. slope of the ground). This means that

equal amounts of positive and negative work are being performed on the body by all

sources. While muscles are the only source of positive work, there are other sources of

negative work in addition to muscle. These include air resistance, damping within the shoe

sole and movement of soft tissue. The first two are known to be small during walking; it is

believed that muscles must perform a substantial fraction of the required negative work.


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Muscles do not act on the environment directly. Instead, muscles act on the body's skeleton

which functions as a system of levers to perform the required power. As a consequence,

positive and negative muscle power is seen externally as positive and negative joint power.

Fig. 1 presents net joint power data, calculated using inverse dynamics for the human knee

measured during walking at a moderate speed (subject mass = 58 kg; speed = 1.3 m/s; step

frequency = 1.8 Hz. . For the angle plot, 180 degrees is full knee extension and knee flexion

is


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a net increase in positive muscle power and metabolic cost. In addition to the power

generating muscle fibres, muscles have elastic elements such as tendons. This provides a

mechanism to store and return elastic energy saving on positive muscle fibre power

production. Regions of negative joint power may actually be times at which elastic energy

is being stored and attempting to harvest energy will increase the total amount of positive

work and metabolic cost. In short, regions of negative joint power are best viewed as

potential regions for energy harvesting and determining their appropriateness requires

experimentation. We focused on the swing phase extension because a) there is a large

amount of negative joint power performed, b) the knee flexors, which act also to extend the

hip, are lengthening because the knee is extending and the hip is flexing suggesting that

they are indeed performing negative work, and c) The energy harvester acts as a rotary

damper element in which the reaction torque is proportional to the angular velocity. This

property favours energy harvesting during the end of swing phase where the angular

velocity is large, allowing efficient power generation with a miniature generator and small

gear ratio gear train.

Fig(2.5).Typical walking mechanics and muscle activity. Subject mass=58 kg; speed=1.3 m/s; step

frequency=1.8 Hz [1]


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2.3.2 Energy Harvesting Methods

In light of the distinct functions of muscle, we distinguish between two general methods of

harvesting energy: parasitic and mutualistic. For parasitic energy harvesting, the electricity

is harvested at the expense of metabolic energy of the user. In this method, the energy is

harvested during the periods when muscles normally perform positive work, causing

muscles to perform more positive work than they would otherwise. On the other hand,

mutualistic energy harvesting is accomplished by selectively harvesting energy at times and

in locations when muscles normally decelerate the body. Rather than braking entirely with

muscles, a generator would perform some of the required negative work converting the

mechanical energy of the body into electrical power. In this manner, mutualistic energy

harvesting would be similar to regenerative braking in hybrid cars.

Fig (2.6).Biomechanical energy harvester.(A) harvester are worn in both legs.(B) mechanical design [2]

It is a challenge, however, to produce substantial electricity from walking. Most energy

harvesting research has focused on generating electricity from the compression of the shoe

sole, with the best devices generating 0.8 W.A noteworthy departure is a spring-loaded

backpack that harnesses the vertical oscillations of a 38-kg load to generate as much as 7.4

W of electricity during fast walking. This device has a markedly low cost of harvesting

(COH),a dimensionless quantity defined as the additional metabolic power in watts required

to generate 1 W of electrical power

(1)


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Where refers to the difference between walking while harvesting energy and walking

while carrying the device but without harvesting energy. The COH for conventional power

generation is simply related to the efficiency with which (i) the device converts mechanical

work to electricity and (ii) muscles convert chemical energy into positive work

(2)

The backpacks device efficiency is about 31%, and muscles peak efficiency is about 25%,

yielding an expected COH of 12.9. But the backpacks actual COH of 4.8 3.0 (mean SD)

is less than 40% of the expected amount. Its economy appears to arise from reducing theenergy expenditure of walking with loads. No device has yet approached the power

generation of the backpack without the need to carry a heavy load.

We propose that a key feature of how humans walk may provide another means of

economical energy harvesting. Muscles cyclically perform positive and negative mechanical

work within each stride. Mechanical work is required to redirect the bodys centre of mass

between steps and simply to move the legs back and forth. Even though the average

mechanical work performed on the body over an entire stride is zero, walking exacts a

metabolic cost because both positive and negative muscle work require metabolic energy.

Coupling a generator to leg motion would generate electricity throughout each cycle,

increasing the load on the muscles during acceleration but assisting them during

deceleration. Although generating electricity during the acceleration phase would exact a

substantial metabolic cost, doing so during the deceleration phase would not, resulting in a

lower COH than for conventional generation. An even lower COH could be achieved by

selectively engaging the generator only during deceleration, similar to how regenerative

braking generates power while decelerating a hybrid car. Here, generative braking

produces electricity without requiring additional positive muscle power. If implemented

effectively, metabolic cost could be about the same as that for normal walking, so energy

would be harvested with no extra user effort.

We developed a wearable, knee-mounted prototype energy harvester to test the generative

braking concept. Although other joints might suffice, we focused on the knee because it


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performs mostly negative work during walking. The harvester comprises an orthopaedic

knee brace configured so that knee motion drives a gear train through a one-way clutch,

transmitting only knee extension motion at speeds suitable for a dc brushless motor that

serves as the generator. For convenient testing, generated electrical power is then dissipated

with a load resistor rather than being used to charge a battery. The device efficiency,

defined as the ratio of the electrical power output to the mechanical power input, was

empirically estimated to be no greater than 63%, yielding an estimated COH for

conventional generation of 6.4. A potentiometer senses knee angle, which is fed back to a

computer controlling a relay switch in series with the load resistor, allowing the electrical

load to be selectively disconnected in real time. For generative braking, we programmed the

harvester to engage only during the end of the swing phase, producing electrical power

while simultaneously assisting the knee flexor muscles in decelerating the knee. We

compared this mode against a continuous-generation mode that harvests energy whenever

the knee is extending. We could also manually disengage the clutch and completely

decouple the gear train and generator from knee motion. This disengaged mode served as a

control condition to estimate the metabolic cost of carrying the harvester mass, independent

of the cost of generating electricity.

Energy-harvesting performance was tested on six male subjects who wore a device on each

leg while walking on a treadmill at 1.5 m/ s .We estimated metabolic cost using a standard

respirometry system and measured the electrical power output of the generator. In the

continuous-generation mode, subjects generated 7.0 0.7 W of electricity with an

insignificant 18 24 W (P = 0.07) increase in metabolic cost over that of the control

condition. In the generative-braking mode, subjects generated 4.8 0.8 W of electricity

with an insignificant 5 21 W increase in metabolic cost as compared with that of the

control condition (P = 0.6). For context, this electricity is sufficient to power 10 typical cell

phones simultaneously. The results demonstrate that substantial electricity could be

generated with minimal increase in user effort.

The corresponding COH values highlight the advantage of generative braking (Fig. 4).

Average COH in generative braking was only 0.7 4.4; less than 1 W of metabolic power

was required to generate 1 W of electricity. This is significantly less than the COH of 6.4

expected for conventional generation (P = 0.01). The COH in continuous generation, 2.3

3.0, was also significantly lower than that for conventional generation (P = 0.01), indicating


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that the former mode also generated some of its electricity from the deceleration of the

knee. The difference between the two modes, 2.2 0.7 W of electricity, came at a

difference in metabolic cost of 13 12 W (P = 0.05). A COH taken from the average ratio

of these differences yields 5.7 6.2, which is nearly the same as that expected of

conventional generation (P = 0.4). This indicates that continuous generation of power at the

knee during walking produces electricity partially by conventional generation with a high

COH and partially by generative braking with a very low COH. But generative braking,

with less than one-eighth the COH of conventional generation, benefits almost entirely from

the deceleration of the knee.

Fig(2.7).Theoretical advantages of generative braking during cyclic motion, comparing the back-and-forthmotion of the knee joint without power generation (A) against a generator operating continuously (B) and

against a generator operating only during braking (C). Each column of plots shows the rate of work performedby muscles (work rate) and the electricity (elect. power) generated over time, as well as the average metabolicpower expended by the human and the resulting average electrical power (ave. power bar graphs). In (B) and(C), work rate is compared against that for (A), denoted by dashed lines, and average power is shown as thedifference ( ave. power) with respect to (A) [3]


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Fig. (2.8) . Biomechanical energy harvester. (A) The device has an aluminium chassis (green) and generator(blue) mounted on a customized orthopedic knee brace (red), totalling 1.6-kg mass, with one worn on each leg.(B) The chassis contains a gear train that converts low velocity and high torque at the knee into high velocityand low torque for the generator, with a one-way roller clutch that allows for selective engagement of the geartrain during knee extension only and no engagement during knee flexion. (C) The schematic diagram shows

how a computer-controlled feedback system determines when to generate power using knee-angle feedback,measured with a potentiometer mounted on the input shaft. Generated power is dissipated in resistors. Rg,generator internal resistance; RL, output load resistance; E(t), generated voltage. [3]

This preliminary demonstration could be improved substantially. We constructed the

prototype for convenient experimentation, leading to a control condition about 20% more

metabolically costly than normal walking: The disengaged clutch mode required an average

metabolic power of 366 63 W as compared with 307 64 W for walking without wearing

the devices. The increase in cost is due mainly to the additional mass and its location,

because the lower a given mass is placed, the more expensive it is to carry. Although the

current increase in metabolic cost is unacceptably high for most practical implementations,

revisions to improve the fit, weight, and efficiency of the device can not only reduce the

cost but also increase the generated electricity. A generator designed specifically for this

application could have lower internal losses and require a smaller, lighter gear train.

Commercially available gear trains can have much lower friction and higher efficiency, in

more compact and lightweight forms. Relocating the device components higher would

decrease the metabolic cost of carrying that mass .A more refined device would also benefit

from a more form-fitting knee brace made out of a more lightweight material such as carbon

fiber.


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Several potential applications are especially suited for generative braking. These include

lighting and communications needs for the quarter of the worlds population who currently

live without electricity supply. Innovative prosthetic knees and ankles use motors to assist

walking, but battery technology limits their power and working time. Energy harvesters

worn on human joints may prove useful for powering the robotic artificial joints. In

implantable devices, such as neurostimulators and drug pumps, battery power limits device

sophistication, and battery replacement requires surgery. A future energy harvester might be

implanted alongside such a device, perhaps in parallel with a muscle, and use generative

braking to provide substantial power indefinitely. Generative braking might then find

practical applications in forms very different from that demonstrated here.

2.3.3 Device DesignThe biomechanics of walking presented four main challenges for designing a device to

harvest energy from the motion of the knee joint. The first challenge was to determine an

effective mechanism for converting biomechanical power into electrical power. This

generator had to be worn on the body so it needed to be small and lightweight. The second

design challenge was to determine a mechanism for converting the knee joint power into a

form suitable for efficient electrical power generation. As described in the previous section,

knee joint power is intermittent, bi-directional and has particular speed and torque

characteristics. The third challenge was to optimize the system parameters in order to

maximize the electrical power generation without adversely affecting the walking motion.

At any given point in the walking cycle, there is only a certain amount of mechanical power

available for harvesting from the kneeattempting to harvest too much power will cause

the user to limp or stop walking while harvesting too little results in less electrical power

generated. The final design challenge was to determine a mechanism for selectively

engaging power generation during swing extension to harvest energy using generative

braking.

To meet these design challenges, our device operated about the knee to take advantage of

the large amount of negative work that muscles perform about this joint (Fig 2). It used a

one-way clutch to transmit only knee extensor motions, a spur gear transmission to amplify

the angular speed, a brushless DC rotary magnetic generator to convert the mechanical

power into electrical power, and a control system to determine when to open and close the


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power generating circuit based on measurements of knee angle. A customized orthopedic

knee brace supported the hardware and distributed the device reaction torque over a large

leg surface area. For convenient experimentation, the control system resided on a desktop

computer and resistors dissipated the generated electrical power. The device was efficient

and the control system was effective at selectively engaging power generation.

Consequently, people were able to generate substantial amounts of electrical power with

little additional effort.

Fig (2.9) The Biomechanical energy harvester comprises an aluminium chassis and generator mounted on

customized knee brace [1]

2.3.4 Mechanical Design Of Harvester

At the end of swing phase, knee angular velocity is less than 100 rpm, and the peak knee

joint torque is around 20 N m (Fig. 1). The energy harvester accepts the input to generate

electrical power and generates enough reaction torque to match the joint torque normally

produced by muscles. The matching of the joint torque and the harvester braking torque is

critical since too much braking torque will interfere the normal walking and too small

torque will not assist the knee flexors enough. The correct torque is achieved by a properly

designed mechanical system. The mechanical system consists of a chassis and a

transmission. The input shaft accepts the knee motion at 1:1 ratio through a single hinge

knee brace. A roller clutch on the input shaft couples the harvester with knee motion during


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knee extension phase, and decouples the harvester from knee motion during the knee

flexion phase.

The normal knee joint mechanical power is computed as:

(1)

Where Pk is the knee joint mechanical power, MK is the knee joint torque from inverse

dynamics; K is the knee angular velocity.

After feeding through the roller clutch, the input angular velocity to the gear train is zero

during knee flexion phases.

The angular velocity is then amplified by the gear train before being applied to the

generator.

(2)

Where rt is the transmission gear ratio. The gear train will spin the generator at a speed of

g and the generator will convert the input mechanical power into electrical power. The

generated voltage is computed by the following equation

(3)

Where Kg is the back electromotive force (EMF) constant which gives the voltage per unit

of rotational velocity. This design parameter of generator depends on the total number of

turns in the armature winding, the number of parallel paths, the number of poles, and the

magnetic flux per pole. A generator with more coils, poles, and stronger flux densitynormally gives a larger Kg. For a motor, the speed constant is the reciprocal of the back

EMF constant.

When connecting a load to the generator, there will be current I in the complete circuit

(4)


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WhereRg is terminal resistance of the generator .Rlis the external load we connect to the

generator. Vlis the output voltage of the generator. The output electrical power is

(5)

The power dissipated by the generator is computed as

(6)

When generating electrical power, the generator produces a

reaction torque that acts on the gear train,

(7)

WhereKm is the torque constant which equals to the back EMF constant.

The reaction torque is amplified by the gear train before being applied to the input shaft and

knee joint. The reaction torque applied to the joint is

(8)

Where t is the efficiency of the gear train. The mechanical power absorbed by the

harvester is the product of the reaction torque and the knee angular velocity,

(9)

The efficiency of the harvester is the ratio between the generated electrical energy and

required mechanical energy,

(10)

For the energy harvester, we want to maximize the electrical power output of (5) and the

mechanical to electrical efficiency of (10) while producing a reaction torque (8) matching

the joint torque normally produced by muscles at the end of swing phase. The design

parameters are gear ratio r, output resistanceRl , the speed constant Kg, and the terminal

resistanceRg. Regardless of the choice of gear ratio and output resistance, a generator with

a smaller speed constant and a smaller terminal resistance will result in higher power output


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and higher efficiency. However, reducing the speed constant and the terminal resistance

means an increase in the weight of the generator. There is a trade-off between the weight

and the preferred generator parameters. We selected a motor with a speed constant of 258

rpm/v, the terminal resistance Rg = 1.03, and a mass of 110 g. After choosing the

generator, we found the output resistance and the gear ratio to maximize the efficiency and

match the reaction torque with the knee joint torque. Through simulation, we found an

optimal combination of gear ratio as 113:1 and output resistance of 5 such that the

electrical power output and mechanical to electrical efficiency are maximized. We

considered the friction of the gear train but neglected the inertia of the transmission and the

generator.

Fig(2.10) Control signals based on knee angle and angular velocity [2]


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2.3.5 Control System Design

In order to achieve generative braking, we designed a control system that selectively

engages/disengages power generation during a gait cycle. The control system consists of a

potentiometer that measures knee angle, an algorithm that generates the control commands,

and an electrical switch that accepts the control command to open/close the circuit between

the generator and resistors. When the circuit is closed, the harvester produces a braking

torque that acts on the knee joint. The control system is implemented on Simulink, compiled

using Real Time Workshop and executed at 1 KHz using Real Time Windows Target on a

desktop computer. This allows for rapid prototyping of the control system. Data acquisition

of the potentiometer data and the control commands to the switch is accomplished by an

A/D and D/A board through the computer.

The potentiometer mounted on the input shaft measures knee joint angle in real time. The

knee joint angle signal is first filtered by a low-pass filter, and then differentiated to get the

angular velocity. The control algorithm uses knee angle and angular velocity to distinguish

different phases of the gait cycle. The logic of the control algorithm is as the following:

(1). If the angular velocity goes across zero upward and the knee angle is small, it is the

start of swing phase knee extension (Point A on Fig.2.10).

(2). If the angular velocity goes across zero upward and the knee angle is large, it is that

start of stance phase knee extension (Point C on Fig. 2.10).

(3). If the angular velocity goes across zero downward between stance phase knee extension

and swing phase knee extension, it is the start of the pre-swing knee flexion (Point B on Fig.

2.10).

(4). If the angular velocity goes across zero downward between swing phase knee extension

and stance phase knee extension, it is the start of stance phase knee flexion (Point D on Fig.

2.10).

Since we target the end of swing phase to harvest electrical power, the power generation

engagement signal is generated by adding a 70-90 ms delay to the detected start of swing

phase knee extension, which is approximately when knee flexor muscles normally become

active to brake knee extension. The disengagement signal is generated by adding a delay to

the start of the stance phase flexion. Instead of turning off energy harvesting at the


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beginning of the stance phase, we keep the harvester on during the stance flexion phase to

allow the generator to harvest the kinetic energy remaining in the transmission and the

generator inertia from the swing phase knee extension. This does not generate extra

resistance for the stance flexion phase since the harvester is decoupled from knee motion

during knee flexion by the roller clutch.

An example result of the control system is shown in Figure. The results indicate that the

control system effectively engages power generation at the middle of the swing extension

phase and disengages at the end of stance flexion phase. Human subject testing

demonstrates that the control system is robust with respect to the variation in knee profile

between subjects. The control system correctly engaged/disengaged power generation for

over 50,000 gait cycles without a failure.

2.3.6 Experimental ResultsWe operated the harvester in there modes: disengaged mode, continuous generation mode

and generative braking mode. In disengaged mode, the roller clutch is manually disengaged

so that the transmission is never in motion. This mode serves as a control condition for

human subject experiments to account for any physiological changes that result fromcarrying the added mass independent of physiological changes resulting from energy

harvesting. In the continuous generation mode, energy harvesting is not selective as the

power generation circuit is always completed. In the generative braking mode the control

system selectively engages and disengages power generation to target the negative work

region at the end of walking swing phase.

Ergometer testing served two purposes. One was to evaluate the harvester efficiency of

converting mechanical power to electrical power, which was used later to determine the

relationship between the amount of generated electrical power and the metabolic cost. The

other purpose was to determine the amount of braking torque produced by the harvester.


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Fig(2.11) Test ergometer for the efficiency evaluation [2]

We designed a test ergometer to drive the harvester with a specified kinetic profile. The

kinetic profile was set as the average knee angle measured during our human subject trials.

By measuring the angular velocity, reaction torque and electrical power generation, we

calculated the efficiency as the ratio between the generated electrical power and the input

mechanical power. To determine the reaction torque produced by friction, the inertia of the

generator and gear train, we performed a test under an open switch condition where no

electrical power was generated. Typical measurements of velocity, torque, computed

mechanical power, and measured electrical power in generative braking, continuous

generation, and open switch modes are shown in figure. The harvester has an efficiency of

63% in continuous generation mode and 56% in generative braking mode. The efficiency in

generative braking mode is lower because the harvester spends a greater amount of time

dissipating mechanical energy without producing electrical power. To determine the

sensitivity of the calculated efficiency to the variation of knee kinematics, we scaled the

input angular velocity profile by 10% and found only small changes in the efficiency (

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