December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 1

BIOMECHANICS OF MECHANICAL HEART VALVE 1

Benjamín González, Humberto Benítez, Kenneth Rufino, Merisabeth Fernández and Waleska Echevarría 2

Abstract - Heart valves all are prone to disease and

malfunction, and can be replaced by prosthetic heart

valves. The two main types of prosthetic heart valves are

mechanical and bioprosthetic. The mechanical valve is

excellent in terms of durability, but is hindered by its

tendency to coagulate the blood. Bioprosthetic valve is

less durable and must be replaced periodically. All

valve types must be durable, because the body is an

extremely hostile environment for a foreign object,

including prosthetic heart valves. Today, engineers are

researching new designs of prosthetic heart valves. They

use the mechanical properties to make an artificial heart

valve design. An artificial mitral valve is an option for

humans with irreparable valve disease.

Key words -- Heart valve, Biocompatibility, Alumina,

Titanium, Biomaterial, Polyether urethane, Polyester,

Pyrolitic carbon.

INTRODUCTION

Heart valves prevent the backflow of blood, which ensures

the proper direction of blood flow through the circulatory

system. Without these valves, the heart would have to work

much harder to push blood into adjacent chambers. The

heart is composed of 4 valves (Figure 1). The Tricuspid

valve is between the right atrium and right ventricle. The

Pulmonary valve is between the right ventricle and the

pulmonary artery. The Aortic valve is between ventricle

and the aorta and the Mitral valve is between the left

atrium and left ventricle. It opens and closes to control

blood flowing into the left side of the heart.

Heart valves open like a trapdoor. The leaflets of the mitral

valve open when the left atrium contracts, forcing blood

through the leaflets and into the left ventricle. When the

left atrium relaxes between heart contractions, the flaps

shut to prevent blood, that has just passed into the left

ventricle, from flowing backward.

1 This review article was prepared on December 8,

2003 for the course on Mechanics of Materials - I.

Course Instructor: Dr. Megh R. Goyal. Professor in

Biomedical Engineering, General Engineering

Department, PO BOX 5984, Mayagüez Puerto Rico

00687-5984. For details contact:

m_goyal@ece.uprm.edu or visit at:

http://www.ece.uprm.edu/m~goyal/home.htm

2 The authors are in the alphabetical order.

The mitral valve, which lies between the two left chambers

of the heart, consists of two triangular-shaped flaps of

tissue called leaflets. The leaflets of the mitral valve are

connected to the heart muscle through a ring called the

annulus, which acts like a hinge.

The mitral valve is anchored to the left ventricle by

tendonlike cords, resembling the strings of a parachute,

called chordae tendineae cordis.

When working properly, heart valves open and close fully.

In mitral regurgitation, the mitral valve does not open or

close properly. Some blood from the left ventricle flows

backward into the left atrium with each heartbeat.

Regurgitation refers to the leakage (backflow) of blood

through a heart valve.

Figure 1. Heart Valves [1].

Heart Valve Problems [7]

There are numerous complications and diseases of the heart

valves that can prevent the proper flow of blood. Heart

valve diseases fall into two categories: Stenosis and

Incompetence. The stenotic heart valve prevents the valve

from opening fully, due to stiffened valve tissue. Hence,

there is more work required to push blood through the

valve. Whereas, the incompetent valves cause inefficient

blood circulation and cause backflow of blood in the heart,

called as regurgitation.

Treatment Options [22]

On a large scale, medication is the best alternative, but in

some cases defective valves have to be replaced with a

prosthetic valve in order for the patient to live a normal

life. An enormous amount of research and development

has proven to be most beneficial, as prosthetic heart valve

technology has saved thousands of lives. Engineers and

The best place to find helping hand is at the end of your arm --- Swedish Proverb.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 2

scientists have done much work to design a valve that can

withstand millions, if not billions, of cardiac cycles. The

two main prosthetic valve designs include mechanical and

bioprosthetic (tissue) Heart Valves.

Mitral Valve Replacement [22]

Valve replacement is done when valve repair is not

possible. Artificial Heart valve is the last solution for

people with a damage heart valve caused by any disease as

regurgitation, etc… In valve replacement surgery, an

artificial prosthetic valve replaces the damaged mitral

valve. The two types of artificial valves are mechanical and

tissue. Mechanical valves, which are made of biomaterials,

may last a long time. However the patient with a

mechanical valve must use an anticoagulant medication

such as warfarin (Coumadin, Panwarfin) for the rest of life

to prevent blood clots from forming on the valve. If a blood

clot forms on the valve, the valve won’t work properly. If a

clot escapes the valve, it could lodge in an artery to the

brain, blocking blood flow to the brain and causing a

stroke. Tissue valves are made of biological tissue such as a

pig’s valve. These kinds of valves are called bioprostheses.

These may wear out over time and may need to be replaced

in another operation. However the tissue valve can avoid

use of long-term anticoagulation medication.

Mitral valve repair or replacement involves open-heart

surgery. Through an incision in the breastbone (sternum),

the heart is exposed and connected to a heart-lung machine

that assumes the breathing and blood circulation during the

procedure. The surgeon then replaces or repairs the valve.

After the operation, which lasts several hours, the patient

spends one or more days in an intensive care unit, where

the general recovery is closely monitored.

History and Advances of Artificial Heart Valves [1]

The first mechanical prosthetic heart valve was implanted

in 1952. Over the years, 30 different mechanical designs

have originated worldwide. These valves have progressed

from simple caged ball valves, to modern bileaflet valves.

The caged ball design is one of the early mechanical heart

valves that use a small ball that is held in place by a welded

metal cage. The ball in cage design was modeled after ball

valves used in industry to avoid backflow. Natural heart

valves allow blood to flow straight through the center of

the valve. This property is known as central flow, which

keeps the amount of work done by the heart to a minimum.

With non-central flow, the heart must work harder to

compensate for the momentum lost due to the change of

direction of the fluid. Caged-ball valves completely block

central flow; therefore the blood requires more energy to

flow around the central ball. In addition, the ball may cause

damage to blood cells due to collision. Damaged blood

cells release blood-clotting ingredients; hence the patients

are required to take lifelong prescriptions of anticoagulants.

For a decade and a half, the caged ball valve was the best

artificial valve design. In the mid-1960s, new classes of

prosthetic valves were designed that used a tilting disc to

better mimic the natural patterns of blood flow. The tilting-

disc valves have a polymer disc held in place by two

welded struts. The disc floats between the two struts in

such a way, as to close when the blood begins to travel

backward and then reopens when blood begins to travel

forward again. The tilting-disc valves are vastly superior to

the ball-cage design. The titling-disc valves open at an

angle of 60° and close shut completely at a rate of 70

times/minute. This tilting pattern provides improved central

flow while still preventing backflow. The tilting-disc valves

reduce mechanical damage to blood cells. This improved

flow pattern reduced blood clotting and infection.

However, the only problem with this design was its

tendency for the outlet struts to fracture as a result of

fatigue from the repeated ramming of the struts by the disc.

Bileaflet valves were introduced in 1979. The leaflets

swing open completely, parallel to the direction of the

blood flow. The bileaflet valves were not ideal valves.

The bileaflet valve constitutes the majority of modern valve

designs. These valves are distinguished mainly for

providing the closest approximation to central flow

achieved in a natural heart valve.

Mechanical Heart Valve

Prosthetic Heart Valves are fabricated of different

biomaterials. Biomaterials are designed to fit the peculiar

requirements of blood flow through the specific chambers

of the heart, with emphasis on producing more central flow

and reducing blood clots. Some of these biomaterials are

alumina, titanium, carbon, polyester, polyurethane etc…

The mechanical properties of these biomaterials involve

how a material responds to the application of a force. The

three fundamental types of forces that can be applied are

stretching (tension), bending, or twisting. Materials

respond to the forces by deforming (changing shape). An

elastic response is reversible, while an inelastic response is

irreversible. In the elastic region, an elastic modulus relates

the relative deformation a material undergoes to the stress

that is applied. The transition between elastic deformation

and failure occurs at the yield point (or stress) of the

material. In designing a component with the material, an

inelastic response is considered failure. Failure can be

plastic deformation or ductile failure. It can also be

breaking, including brittle failure or fracture. Mechanical

properties of a material in the range of elastic behavior

include its elastic modulus under tension and shear stresses,

its Poisson’s ratio, its resilience, and its flexural modulus.

The transition to failure is denoted by the yield stress or

breaking strength of the material.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 3

BIOMATERIALS

The requirements for an artificial heart valve are

staggering. The valve must be easy to insert. It must last a

long time. It must be able to open and close 35 million

times a year for 20 to 50 years. It must allow high blood

flow with minimal turbulence and must not leak. The valve

also should not cause blood clots and also:

� Collapse to 5 mm when crimped.

� Top of stent expands to 25 mm.

� Middle of stent expands to 30 mm.

� Bottom of stent expands to 25 mm.

� Deployed height is 25.4 mm.

� Collapse to �5 mm when crimped.

� Barrel shaped.

� Top stent expands to 30 mm.

� No damage to leaflets.

� Length is 12.7 mm - 25.4 mm.

1. Alumina (Al2O3) : aluminium oxide [17]

Alumina (Al2O3) is a bioinert material. Bioinert materials

do not chemically react with the local chemicals As a

result, cells can survive next to the material but do not

form a union with it. Often fibrous protective cells grow

near the implant surface to protect local cells from

mechanical damage. Bioinert materials were first used for

prosthetics. These materials can be very strong but have

the disadvantage of not bonding to the local cells.

Numerous problems have been encountered in anchoring

the bioinert implants to bone. In early implants, some

implants became deformed or displaced, causing serious

damage to the surrounding tissue.

Alumina is a traditional ceramics that offer many

advantages compared to other biomaterials. These are

harder and stiffer than steel; more heat and corrosion

resistant than metals or polymers; less dense than most

metals and their alloys; and their raw materials are both

plentiful and inexpensive. Design requirements for alumina

as a biomaterial are:

� High fluid resistance.

� Avoid hemorrhage.

� Low incident of thromboembolism.

� Be economic.

� High performance.

� Avoid stiffening of the leaflets.

� Optimal designs.

� Good thermal conductivity.

� Ability to open and close 35 million times a year

for 20-50 years.

� Biocompatibility.

� Avoid blood clots.

� Available easily.

Alumina as biomaterial

Alumina is the most widely and versatile ceramic. Much of

the research on this ceramic was done during the 1950s and

1960s. Alumina is chemically stable against most

environments except hydrofluoric acid and some molten

salts. These traditional ceramics set upon hydration if

produced in the special form of re-hydratable alumina

cement (more commonly in the form of calcium aluminate

cement). Also alumina is widely used for medical implants

like mechanical heart valve. This type of ceramic is also

used in several medical fields as dentistry, orthopedical and

cardiologist application.

It is a ceramic, non-metallic, and inorganic compound that

displays great strength and stresses resistance to corrosion

wear and low density. Alumina is a highly bioinert

material and resistant to most corrosive environments,

including the highly dynamic environment of the human

body.

Compatibility between Bioceramics and the Human

Environment [17 and 26]

The major problem on implants designs is the fairly limited

choice of materials, and consequently, determination of the

compatibility of the material of choice with the tissue.

There are no standard methods of compatibility testing and

the number of variables involved is usually much larger

than the typical engineering problem. For example, human

blood is 1/3 as salty as seawater, stays at a steady 37ºC, and

contains active enzymes (the immune system).

Human body is one of the most corrosive environments that

inorganic substances can encounter. Furthermore, as the

various metabolic processes occur in an organism the

various complex molecules that may enclose a substance

continually change in concentration and variety. Lactic acid

produced from muscle cells during anaerobic cellular

respiration is a prime example. Additionally, the time

factor of the “compatibility reaction” is important; the

implant - tissue interaction is a sequential chain of

reactions, characteristic for the material and the patient.

Beyond the chemical factors, the response of tissue

depends additionally on geometric characteristics of the

implant, e.g. shape, size, surface/volume ratio. These

factors will generally determine the state of stress at the

interface and thus could interfere with the interfacial

reactions. Porosity and its size distribution within an

implant have been shown to affect the interactions. It has

been generally established, that tissue will grow into pores

larger than ~120 nm.

Most polymers seem to be slowly “digested” by the human

body and metals are slowly corroded: high concentration of

metallic elements has been detected close to the (metallic)

implant surface. Passive oxide film can significantly slow

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 4

down the reactions with tissue, rendering titanium or

stainless steel virtually neutral. However, polymeric and

metallic implants are generally classified as “temporary”

implants, with very low adhesive strength of attachment to

the tissue.

Alumina is considered bioinert due to a thin layer of

titanium ions on its surface, although some studies show

that the body can absorb alumina. Various researchers have

found alarmingly low levels of Al in rats’ nervous systems

after the 20-week postoperative period.

The bioinert ceramics, like ZrO2, Al2O3, SiC, Si3N4 do not

develop strong interfaces, but also do not liberate ions into

the internal environment. This is at the expense of lesser

mechanical performance and reliability of ceramics, as

compared to metals. A compromise is ceramic-coated

metal, although some additional liabilities are created (e.g.

adhesion of the coating; increased processing costs etc.).

Alumina Mitral Valve [26]

Ceramic materials are somewhat limited in applicability by

their mechanical properties, which in many respects are

inferior to those of metals. The principal disadvantage is a

disposition to catastrophic fracture in a brittle manner with

very little energy absorption.

The ceramic mitral valve is comprised of a single crystal

alumina disc and titanium valve ring. Alumina consists of

aluminum and oxygen ions. These ions combine firmly by

ionic bond and are arranged in hexagonal closed packed

structure. The single crystal alumina disk is 1.0 mm thick.

Both mechanical and chemical polishing smoothed the

surfaces. The valve ring was milled from a single piece of

titanium and was coated with Tin by reactive ion plating

(See appendix IV). Alumina has a good blood

compatibility, excellent wear resistance, largely inert and

durability. Alumina mitral valve avoids thrombus

formation and thromboembolism.

Tensile strength of single alumina is more than three times

greater that LTI carbon. Alumina has hardness eight times

greater than LTI carbon. Alumina is insoluble in water and

has high corrosion.

Properties of Alumina

a. Scratch resistance: The extreme hardness of alumina is

second only to a diamond. Metal-on-metal articulations can

be scratched causing an abrasive surface. Foreign debris in

the joint may also accelerate implant wear. Alumina is

more scratch resistant than metal or polyethylene; so it is

most durable than other valve materials.

b. Ion release: Since ceramics do not release ions, there are

no long-term unknowns pertaining to systemic effects due

to ion release with this hard bearing couple, unlike metal-

metal bearing couples.

c. Friction and wetability: A material that holds

lubrication to its surface is considered wet able. Alumina

ceramic is a more wet able material than metal. Lubrication

helps to reduce friction between components. Alumina

ceramic has improved since the 1974. Third generation

materials have nearly twice the strength as the original

material because of enhancements in purity, density and

grain size.

d. Fracture: This property continues to be the primary

concern regarding ceramic components. Improper handling

and implantation, poor implant design and material, or

mismatched components caused fractures in early ceramic

designs. When correctly implanted, the fracture rate has

been reported between one-tenth to one-twentieth of a

percent (0.001 - 0.0005) and it is projected that

contemporary materials will be even lower. Alumina

should be use, only in compression.

e. Strength: Though ceramics are brittle in nature, alumina

ceramic inserts are extremely strong and exceed FDA

Guidance Document standard for ceramic heads of 46 kN

or 10,340 pounds burst strength. This exceeds the strength

of the ceramic head as well as the neck of the femoral stem.

As with any modular interface under load, there is a

potential for micro motion and associated fretting and/or

corrosion. However, the alumina design minimizes the

amount of motion at the taper interface, which should

reduce the corrosion potential.

Alumina mechanical properties are summarized as:

• Good mechanical strength (Figure 3).

• Good thermal conductivity.

• High electrical resistivity.

• High hardness (Figure 2).

• Wear resistant.

• Good chemical stability.

• Largely inert.

• Excellent tribological characteristics.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 5

Figure 2. Hardness of different biomaterials [26].

Valve Problems with Alumina

a. It is a hard material so that machining is difficult.

Therefore, some molding process must be developed which

can produce a finished valve with the accurate internal

shape required to achieve good homodynamic performance.

b. The tissue covering requires a porous, textured alumina

surface on which to anchor itself firmly, but the main body

of the conduit valve must be in the most dense form of

alumina, with virtually no porosity, in order to maintain

structural strength. The molding process must therefore

accommodate variable porosity in some way.

c. Alumina cannot avoid thromboembolism totally.

Alumina Future

Research is being done to combine alumina with other

materials for better heart valve implants. These

experiments try to avoid tromboembolism, which is the

major problem on heart valve implants. Today, there are no

materials to avoid totally thrombosis. Alumina is a

material with good mechanical properties, but mechanical

heart valve manufactures didn’t use for this purpose.

Instead alumina is utilized in dental implants.

Lawsuits against medical device manufacturers,

restructuring of FDA approval procedures, patient

expectations, and the health care reform movement are

changing the future of the medical device community and

shaping the direction of biomaterials research. As a result,

new materials and manufacturers will be required to meet

FDA standards. Another important issue not often

discussed is that implant recipients expect an implant to

function and to last forever.

Biomaterials and implant research will continue to

concentrate on serving the needs of medical device

manufacturers and recipients, as well as medical

professionals, and on developing technologies to meet

those needs. Future biomaterials like alumina will

incorporate biological factors directly into an implant’s

surface to improve biocompatibility and bioactivity. New

projects will be directed at materials development for

improved mechanical integrity, corrosion resistance, and

biocompatibility. Institute engineers will also apply

statistical finite element analysis, stereo imaging strain

analysis, and composite materials to the biomaterials

program. Therefore, alumina is one of these experimental

materials for the future.

Figure 3. Tensile Strength (MPa) of biomaterials [26].

2. Polyester

What is polyester? [27]

Polyester is a synthetic resin formed by the condensation of

polyhydric alcohols with diabasic acids. Polyesters are

thermosetting plastics used in making sythentic fibres and

constructional plastics. It is an extremely resilient fibre that

is smooth, crisp and particularly springy. Its shape is

determined by heat and it is insensitive to moisture. It is

lightweight, strong and resistant to creasing, shrinking,

stretching, mildew and abrasion. It is readily washable and

is not damaged by sunlight or weather and is resistant to

moths and mildew. The following requirements must be

satisfied to use polyester as a biomaterial in heart valves

implants:

• The body’s immune system must not attack the

biomaterial.

• Compatible with body tissues and fluids

• Must have strength, flexibility and hardness

• Must be nontoxic, nonreactive or biodegradable

• The replacement valve must be smooth to prevent

the destruction of blood vessels.

• The valve must also be anchored to the inside of

the heart.

• Must be an elastomer so it can be flexible during

the pumping cycle.

0

200

400

600

800

1000

Alu

min

a

Sta

inle

ss

Ste

el

Titaniu

m

LT

I

Carb

on

0 5 10 15 20

Alumina

LTI Carbon

Titanium

Titanium Nitride

Series1

Series2

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 6

• The material must not exhibit mechanical fatigue

over the device’s lifetime.

• The material’s surface must have an acceptable

low propensity for thrombus formation, as well as

the best possible blood compatibility

• Must not be prone to calcification.

• The material must be easily formed into complex

shapes.

Structure and Physical Properties of Polyester

• Polyester chains tend to be flexible and are easily

entangled or folded.

• Degree of crystallinity is the amount of ordering

in a polymer.

• Stretching or extruding a polymer can increase

crystallinity.

• Degree of crystallinity is also determined by

average molecular mass.

• Bonds formed between polyester chains make the

polyester stiffer.

Development of Polyester in Vascular Surgery

Vascular prostheses fabricated as polyester textile tubes are

most frequently used devices in peripheral vascular surgery

for the replacement of large and medium sized vessels.

Long term results representing a period of follow-up over

15 – 20 years have shown satisfactory results when Dacron

grafts are installed in the aortic and iliac sites. Technical

developments to improve the device over the years have

passed through different generations of concepts. The

relative merits of these different designs are still a matter of

intensive research.

Woven or Knitted Design

Weslowski was the first to recognize the importance of the

porosity within the graft wall for the healing process of the

graft. By using a more open textile structure with large

pores between the polyester fibers it was predicted that

cellular elements and fibrous tissue would be able to

penetrate the interstices of the graft wall and generate a

well-attached, more completely healed surrounding

capsule. Unfortunately, measurements of water

permeability were mistakenly assumed to measure the

porosity of the graft wall, and as a result manufacturers

produced thinner and thinner textile structures using finer

and finer polyester yarns with a view to improve the

healing performance of the prostheses. The creation of the

ultra-light-weight design provided the surgeon with a more

flexible graft that was easier to handle and suture. But too

high water permeability posed difficulties in preclothing

the graft so as to achieve hemostasis. Problems of

hemorrhage at the time of implantations and complications

associated with secondary hematomas around the grafts as

well as prosthetic dilatation and failure caused this concept

to be abandoned in the late 1970s. Surprisingly enough, the

ideal values for the porosity and water permeability of a

vascular prosthesis are defined poorly.

In spite of the success of expanded PTFE grafts that remain

patent for many years without any tissue encapsulation, it is

still believed that complete healing of the luminal surface is

a critically important requirement for long-term patency.

One approach to achieve this was proposed by DeBakey

and to improve the anchoring of fibrous tissue, increase

cellular adhesion, and hence promote the formation

neointima by the use of the velour design. This involves

weaving or knitting rather than straight fibers to give a

rougher, randomized, and more open appearance to the

external and/or internal surface of the graft. External velour

enables better incorporation of the graft within the host

tissue, whereas internal velour encourages the formation of

thicker neointima. This may be less important in clinical

practice since complete endothelialization is never

accomplished in humans. Even so, internal external and

double velour grafts are widely available.

As a result of the complications such as dilatation

associated with the light-weight weft-knitted design,

manufacturers have taken steps to increase the strength of

grafts by using thicker polyester yarns and tighter, more

compact woven constructions. The more open woven

velour constructions should be anastomosed with a larger

than normal bite or cut with a hot cautery in order to reduce

the risk of fraying at the suture line. The regular woven and

low porosity woven design are used widely for the

replacement of the thoracic aorta and for interventions

involving a cardiopulmonary bypass with heparinization.

For those surgeons who prefer the ease of handling and

suturing of the knitted construction, most major models

with the more dimensionally stable warp knitted prosthesis

have now replaced former weft knitted models, thus

ensuring the same good anchorage of the neointima but

avoiding the complications associated with dilatation and

raveling of the textile structure in vivo. The importance of

maintaining the initial strength of vascular prostheses at an

acceptable level is now widely accepted.

Externally Supported Design

The problem of flattening and occlusion of a vascular graft

at the point where it crosses a knee or hip joint is well

known. External reinforcement of the graft by means of a

rigid spiral support has proven to be effective in alleviating

this problem and has found merit in the axillofemoral

position as well. The performance observed during animal

trials as well as clinical observations of explanted devices

suggest that high levels of friction and fatigue can occur to

the textile structure underneath the rigid external support.

This is particularly problematic with those models where

the external support is not well attached to the outer surface

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 7

of the prosthesis. As a result, perhaps the most valuable

application for this type of design is in the axillofemoral

bypass where compression of the graft may occur, when the

patient is lying on the relevant side.

Composite Design

In order to improve the biocompatibility of porous

synthetic grafts, it has been proposed that the polyester

textile structure be impregnated or coated with a

crosslinked protein. A number of different proteins have

been studied, including albumin, collagen, gelatin, elastin,

and chitosan. Observations from comparative in vivo

studies as well as from our explant retrieval program

indicate that the healing process is virtually identical with

the coated and uncoated grafts. The only difference, if any,

resides in the rate of the healing process, wich is slowed by

the addition of the protein coating. Because the prosthesis

is already nonpermeable to blood and ready to implant the

moment it is removed from is sterile packaging, the need

for blood transfusion and preoperative manipulation is

reduced. In addition, since the coating is resorbed slowly, it

has been proposed that antibiotics and growth promoting

factors be added to the protein in order to reduce the risk of

infection and enhance the healing of the neointima.

The cellular seeding of vascular prostheses with endothelial

cells appears to be a very promising technique. The

experimental research to date has improved our

understanding of the different functions of the endothelial

cell and its interactions with blood. However, the efficiency

of the cell seeding procedure leaves much to be desired,

and, while the technique has proven useful in a few human

trials, it is not yet ready for routine clinical use. In the long

term, there it do appear to be beneficial in using this

technology, particulary in femoropopliteal and distal sites

where the rate of flow is limited and in reducing the

incidence of infection.

Another type of surface coating proposed for vascular

grafts involves the use of the plasma graft polymerization

process. This technique can modify the surface chemistry

and hence the biocompatibility of a synthetic material.

Typically, plasma of fluorethylene gas is generated in a

evacuated chamber containing the prosthesis by means of a

high-frequency magnetic field. The free radicals so

produced react rapidly with each other and with any surface

they encounter, depositing a thin layer of a fluorocarbon

polymer on the polyester fibers of the vascular graft. The

flow surface is thus likely to be more hydrophobic and

biocompatible. Preliminary results in animals have so far

been promising, but they have not been confirmed in

humans.

Biocompatibility

Although saphenous vein remains the material of choice

for vascular reconstruction, fem-popliteal grafts composed

of polyester have gained in popularity due to the similar

rates of occlusion and ease of use. Despite these

advantages, there is continuing concern over the

nevertheless high rates of occlusion, for both native and

synthetic grafts which significantly contributes to the

greater than 30% failure rate experienced within one year.

As these failures often lead to limb loss and other serious

complications, the availability of a non-thrombogenic graft

or one with reduced thrombo-genicity would have

significant clinical impact and could serve to reduce the

incidence of thrombosis related complications. Polymeric

biomaterial surfaces such as polyester and PTFE are

intrinsically thrombogenic. Grafts composed of these

materials can be surface modified in order to reduce this

inherent thrombogenicity. Heparin treatment of the surfaces

of a number of medical devices such as catheters, heart

valves, stents and bypass circuitry has successfully been

used as a means of reducing surface thrombogenicity.

In 1991, InterVascular developed the concept of adding

unfractionated high molecular weight heparin to the inner

lumen of a graft through a stable bonding process. It was

believed that this modification of the graft surface could

significantly reduce its thrombogenicity and potentially

improve graft performance and clinical outcome. Heparin is

coupled to the InterGard Heparin surface using tri-

dodecylammonium chloride (TDMAC) which forms an

insoluble complex with heparin and in turn binds with high

affinity to the polyester flow surface through its long

hydrophobic tails. The heparinized graft is then coated with

collagen which acts as a barrier to prevent rapid release of

the heparin from the graft surface. A series of studies was

performed to evaluate the safety and efficacy of the

InterGard heparin coated graft. Animal studies were

performed to confirm that no bleeding complications and

good healing characteristics were associated with the use of

the heparin bonded graft.

In addition, complete ISO 10993 biocompatibility tests

were performed to assure that safety and biocompatibility

requirements were met. Bench studies were conducted to

evaluate the retention of heparin on the InterGard heparin

graft in a simulated model of circulation using

physiological flow rates and pressures. In these studies,

heparin levels remained constant for 7 days in the

InterGard heparin collagen coated graft but declined

dramatically in the non-collagen coated graft demonstrating

that the stable bonding process of ionic coupling to

TDMAC followed by hydrophobic interaction with

polyester immobilizes the heparin to the graft. Furthermore,

the collagen coating helps retain the heparin complex

preventing its premature release. Additional studies were

performed which demonstrated that the heparin-collagen

coating dramatically reduces the deposition of fibrin (a

measure of thrombogenicity) relative to uncoated polyester

grafts. These studies coupled with on going promising

clinical data continue to support the safety, utility and

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 8

clinical benefits associated with the InterGard Heparin

graft.

Polyesters Chemical Properties [20]

Polyesters are formed either by a reaction between a dibasic

acid and a dihydroxy alcohol or by the polymerization of a

hydroxy carboxylic acid. The chemical structure of a

polyester is shown in figure 4. Polyesters are naturally clear

and colorless; however they can be colored and made

according to specifications. Polyesters do not show wear

with exposure to poor weather conditions. They are highly

resistant to chemical deterioration, withstanding most

solvents, acids, and salts. They are also resistant to heat

damage and can be made to be self-extinguishing.

Not to be outdone, DuPont was also at the forefront of

polyurethane technology in the U.S., receiving patents in

1942 covering the reactions of diisocyanates with glycols,

diamines, polyesters and certain other active

hydrogencontaining chemicals. From these humble

beginnings emerged the polyurethanes, the most versatile

polymers in the biomaterials armamentarium.

Figure 4. Chemical Structure of polyester [20].

3. Polyurethanes [25 and 27]

Polymers are considered some of the most promising class

of biomaterial. They can be selected according to certain

characteristics such as mechanical resistance, degradability,

permeability, solubility, as well as transparency.

Polyurethanes are the polymers most widely used in the

construction of blood-contacting products and devices.

History of polyurethane

Nineteen eighty-seven marked the 50th anniversary of the

introduction of polyurethanes. Professor Otto Bayer was

synthesizing polymer fibers to complete with nylon when

he developed the first fiber-forming polyurethane in 1937.

Polyurethanes technology

Current activities of suppliers, designers, manufacturers

and physicians clearly indicate that devices manufactured

from synthetic polymers have become an integral part of

health-care technology. Initially focused on life-threatening

situations, their clinical uses now include permanent

implantation (e. g. artificial hearts, hip prostheses,

intraocular lenses), intermediate applications (e. g. contact

lenses, removable dental prostheses, renal dialyzers), and

transient applications (e. g. cardiopulmonary bypass, over-

the-needle catheters, diagnostic and therapeutic catheters).

The polymers used most often in these applications are the

silicone elastomers, the acrylics, polyvinyl chloride,

fluorinated ethylene propylene and polycarbonates.

In the past ten years, research work on the artificial heart

has stimulated interest in this new family of polymers, the

segmented polyurethane elastomers. Originally developed

for commercial applications, these polymers exhibit high

flexure endurance, high strength, and inherent

nonthrombogenic characteristics, and are expected to have

a positive effect on future medical applications. Segmented

polyurethane polymers are widely used as artificial heart,

vascular grafts, catheter, diaphragm of blood pump,

pacemakers wire insulation, heart valves, cardiac-assist

devices, components of hemodialysis units, skin grafts and

blood filters. Since the segmented polyurethane exhibit

high strength, nonthrombogenic characteristics, the most

important applications appear to be in the cardiovascular

area. Because of higher hydrolytic resistance and better

properties at low temperatures, the structures of

polyurethanes prepared from lactones can be used as

medical, solvent-activated, pressure-sensitive adhesives.

Future scope of polyurethane

A material used in the leaflet heart valves, mechanical heart

valve coatings and total artificial heart is polyether-based

polyurethane. However, one drawback of this material is

the absorption of the proteins and thus, the onset of

thrombosis and bacterial infection. The right materials have

the good mechanical properties of polyurethane while

eliminating the risk of thrombosis and bacterial infection.

Unfortunately, scientists have been unable to find a suitable

substitute with such mechanical properties as well as

relative biocompatibility. Therefore, scientists have begun

searching for possible improvements to polyurethane in an

attempt to increase its biocompatibility.

One possible solution to the compatibility problem is to

synthesize a polymer alloy consisting of polyurethane along

with a phospholipid polymer. A current polymer alloy that

has shown promise in combating the onset of thrombosis as

well as bacterial infection is 2-methacryloyloxethyl

phosphorylcholine (PMEH) with segmented polyurethane.

Research on this alloy has shown a significant decrease in

the amount of proteins absorbed at the blood-suface

interface. In fact, when protein adsorption data was

recorded, the amount of the protein adsorbed on the 2-

methacryloyloxethyl phosphorycholine segmented

polyurethane tubing was only 17% of that adsorbed by

segmented polyurethane tubing alone. In similar attempt,

scientists synthesized an alloy of polyurethane with the

addition of poly (tetramethylammonium) oxide and

methylene diphenylene diisocyanate along with chain

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 9

extenders of 3-trinethylammonium-1,2-propanedioliodide

(TMPI) and 3-dimethylamino-1,2-propanedioliodide

(DMP). This alloy had been previously found to restrict the

onset of thrombosis. In an experiment conducted to

determine the protein attachment rate constant of three

polyurethane alloys as well as pure polyurethane. This

polyurethane, labeled PEU-N, was found to have a higher

attachment constant (.00059 cm/min) than either of the

other polyurethane alloys including a phospholipids

polymer alloy (PEU-G). However, PEU-N did have a lower

adhesion constant than pure polyurethane (PEU-B).

Not all polyurethanes are equally effective in their

biocompatibility properties. Polyurethanes comprise a large

family of materials, with urethane linkage being the only

common characteristic. They have been found to vary in

clinical applications. When implanted in the human body,

polyester-based polyurethanes tend to undergo a rapid

hydrolysis and should be avoided in medical applications.

Due to their quick crystallization, polycaprolactone-based

polyurethanes can be used as medical applications, but only

as pressure-sensitive adhesives. Polybutadiene-based

polyurethanes have been investigated, yet no medical

application has been found to date. Castor oil-based

polyurethanes can be used, but due to their poor tear

resistance, have a very limited use in medical applications

since they are virtually insensitive to hydrolysis, and

therefore are very stable in the physiological environment.

4. Polyether urethane

In the preparation of this type of polymers, polyether-based

glycols are used. If they are cured with aromatic diamines

then their structure-property relationships will be very

similar to those of polyester urethanes. At high NCO/NH2

ratios the excess isocyanate forms biuret branch points.

Thus, an increase in cross-linking causes a reduction in

modulus, elongation, compression set, and tears strength.

The secondary reactions occur to a much less extent than

the primary reactions but their importance must not be

underestimated. Formation of allophanates or biurets is

responsible for some of the cross-linking and branching

and therefore has an important influence on the properties

of the polyurethane product.

Elasthane™ polyether urethane [19]

Elasthane™ polyether urethane is a high-strength, aromatic

thermoplastic with a chemical structure and properties very

similar to Pellethane® 2363 polyetherurethane series, which

has been used to fabricate a large number of implantable

devices, including pacemaker leads and cardiac prosthesis

devices such as artificial hearts, heart valves, intraaortic

balloons, and ventricular assist devices. PTG's Elasthane is

designed for chronically-implanted medical devices and

demonstrates an impressive combination of mechanical

properties and biological compatibility. The Polymer

Technology Group developed Elasthane in response to

Dow's decision to limit Pellethane's use in chronically

implanted medical devices. In developing Elasthane, PTG

invested in the same continuous reactor technology to offer

the only Pellethane substitute with the same high molecular

weight and reduced thermal history as Pellethane. PTG has

rigorous quality control and documentation of the

manufacturing procedures, formula optimization, and

precision feed pumps. Formal validation of Elasthane was

accomplished through rigorous short- and long-term testing

in conjunction with a major academic institution and a

medical device company that has since received approval to

implant the material. A comprehensive FDA Masterfile also

backs Elasthane.

Elasthane™ polyether urethane is a thermoplastic

elastomer formed as the reaction product of a polyol, an

aromatic diisocyanate and a low molecular weight glycol

used as a chain extender. Polytetramethylene oxide

(PTMO) is reacted in the bulk with aromatic isocyanate,

4,4'-methylene bisphenyl diisocyanate (MDI), and chain

extended with 1,4-butanediol.

Application of Elasthane™ polyether urethane

Numerous medical devices and technologies have benefited

from the combination of exceptionally smooth surfaces,

excellent mechanical properties, stability, and good

biocompatibility of Elasthane™ polyether urethane.

Pellethane is currently the polyurethane used for the

tricuspid semilunar valves. Due to its high molecular

weight, valves fabricated from Elasthane have shown to

reduce the degree of calcification. Furthermore, Elasthane

that has been chemically modified with polyethylene oxide

(P) and sulfonate (SO) SMEs showed lower surface platelet

adhesion and thrombus formation, suggesting improved

blood compatibility.

Hydrodynamic evaluation of Pellethane valves showed

minimum pressure drop and very low energy losses

compared with other commercially available valves. It was

also found that in durability tests, prototypes have lasted

for 17 years.

Mechanical Properties of Polyether urethane [10]

At lower hardness levels, practically all elastomeric

materials, including polyurethane elastomers, merely bend

under impact. As conventional elastomers are compounded

up to higher hardness they tend to lose elasticity and crack

under impact. On the other hand, polyurethane elastomers

when at their highest hardness levels have significantly

better impact resistance than almost all plastics.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 10

Such great toughness, combined with the many other

outstanding properties associated with the high hardness

polyurethane leads to many applications in engineering.

(Appendix I gives properties of different kinds of

Polyurethanes).

Biocompatibility of Polyether Urethane

The blood contacting surface of some leaflets hearts valves

are made of polyether urethane (PEU, n = 22). This

polyurethane can be resistant to thrombus formation when

processed into an ultra smooth contacting surface.

Elastomeric polyurethanes are inherently thromboresistant.

Although blood compatibility and nonthrombogenicity are

subject to many complex factors, such as polymer surface

composition, device configuration, and blood-flow

characteristics, they tend to perform well in numerous

device configurations. Their apparent thromboresistance is

thought to reside in polyurethane’s ability to preferentially

absorb serum albumin.

When the biomaterial surface comes into contact with

blood, a protein layer of fibrin results from the

polymerization of fibrinogen. When bacteria interacts with

the surface of a blood-contacting biomaterials it does so

through this adsorbed protein layer. Therefore, bacteria can

easily attach itself to the material surface and cause

infection.

When choosing a material to combat bacterial adhesion, it

is essential that the material limits protein adsorption.

Proteins tend to be negatively polarized and thus

hydrophobic in nature. With this in mind, it is beneficial to

select a material that is similarly polarized, thus likely to

repel the proteins from the biomaterial’s surface, hindering

the protein-surface interaction and protein adsorption by

the surface. By disrupting this adsorption of proteins, the

material is less likely to develop a protein layer and less

likely to promote the development of bacterial growth and

infection.

Surface modification of polyurethane heart valves:

effects on fatigue life and calcification

Polyurethane heart valves can be functionally durable with

minimal calcification, in vitro. In vivo, these characteristics

will depend on the resistance of the polyurethane to

thrombogenesis and biodegradation. Surface modification

may improve the polyurethane in these respects, but may

adversely affect calcification and durability. This study

investigates the effects of surface modifications of two

polyurethane heart valves (PEU and PEUE) on vitro fatigue

and calcification behavior. Modifications included heparin,

taurine or aminosilane. Aminosilane modification of PEUE

valves increased durability compared with PEO

modification. Appropiate surface modification may be

useful to improve blood compatibility of implantable

polyurethanes, and may also be advantageous as regards

fatigue durability of flexing materials in long term

applications.

Polyurethane heart valves: Fatigue failure, Calcification

and Polyurethane Structure

Six flexible-leaflet prosthetic heart valves, fabricated from

a polyether urethane urea (PEUE), underwent long-term

fatigue and calcification testing by Dernacca GM,

Guldransen NJ; Wikinson R; and Wheatley DJ. They

discovered that three valves exceeded 800 million cycles

without failure. Three valves failed at 775, 460, and 544

million cycles, respectively. Calcification was observed

with and without associated failure in regions of high

strain. Comparison with similar valves fabricated from a

polyether urethane (PEU) suggested that the PUE is likely

to fail sooner as a valve leaflet. Localized calcification was

developed in PEUE leaflets at the primary failure site of

PEU leaflets, close to the coaptation region of three

leaflets. The failure mode in PEU valves had the

appearance of abrasion wear associated with calcification.

High strains in the same area may render the PEUE leaflets

vulnerable to calcification. Intrinsic calcification of this

tape, however, is a long-term phenomenon unlikely to

cause early valve failure. Both polymers performed

similarly during static in vitro and in vivo calcification

testing and demonstrated a much lesser degree of

calcification than bioprosthetic types of valve materials.

Polyurethane valves can achieve the durabilities required of

an implantable prosthetic valve, equaling the fatigue life of

currently available bioprosthetic valves.

Polyurethane heart valve durability: Effects of leaflets

thickness and material

The durability of a flexible trileaflet polyurethane valve is

determined by the thickness of its leaflets. Leaflet thickness

is also a major determinant of hydrodynamic function. The

study was conducted by Dernacca GM; Guldransen NJ;

Wikinson R; and Wheatley DJ examined valves (n = 31)

with leaflets made of polyether urethane (PEU, n = 22) or a

polyether urethane urea (PEUE, n = 9), of varying

thickness distributions. The valves were subjected to

accelerated fatigue test at 37ºC and failure was monitored.

Leaflet thickness ranged from 60 to 200µm. PEU leaflet

thickness bore no relationship to durability, which was less

than 400 million cycles. PEUE valves, in contrast,

exceeded 800 million cycles. Durability in PEUE valves

was directly related to leaflet thickness ( r = .93, p < 0.001),

with good durability achieve with median leaflet

thicknesses of approximately 150 µm. Thus polyurethanes

valves can made with good hydrodynamic properties and

with sufficient durability to consider potential clinical use.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 11

New polyurethane heart valve prosthesis: Design,

manufacture and evaluation

In light of the thrombogenicity of mechanical valves and

the limited durability of bioprosthetic valves, alternative

designs and materials are being considered for prosthetic

heart valves. A new tri-leaflet valve, made entirely from

polyurethane, has been developed. The valve comprises

three thin polyurethane leaflets (approximately 100µm

thick suspended from the inside of a flexible polyurethane

frame. The closed leaflet geometry is elliptical in the radial

direction and hyperbolic in the circumferential direction.

Valve leaflets are formed and integrated with their support

frame I a single dip coating operation. The dipping process

consistently gives rise to tolerably uniform leaflet thickness

distributions. In hydrodynamic test, the polyurethane valve

exhibits pressure gradients similar to those for a

bioprosthetic valve (St Jude Bioimplant), and levels of

regurgitation and leakage are considerably less than those

for either a bileaflet mechanical valve (St Jude Medical) or

the bioprosthetic valve. Six out of six consecutively

manufactured polyurethane valves have exceeded the

equivalent of 10 years function without failure in

accelerated fatigue tests. The only failure to date occurred

after the equivalent of approximately 12 years cycling, and

three valves have reached 527 million cycles

(approximately 13 years equivalent).

5. Pyrolytic Carbon

Background

Dr. Jack Bokros and Dr. Vincent Gott [11] discovered

pyrolytic carbon, the premier material for artificial heart

valves at General Atomics (GA). In 1966, Dr. Bokros was

working on pyrolytic carbon coatings for nuclear fuel

particles for the GA gas-cooled nuclear power reactors. He

stumbled upon its potential for medical uses through what

has been called “a lesson in serendipity”. He read an article

by Dr. Vincent Gott, who has been testing carbon-based

paint as a blood compatible coating for artificial heart

components. Bokros contacted Gott who initiated the

collaboration.

Dr. Gott was searching for a material to use in artificial

heart valves that did not provoke blood clots and had the

mechanical durability to endure for a recipient’s lifetime.

Pyrolytic carbon, from GA, met both of his need. GA

initiated a development project headed by Dr. Brokros to

add the needed durability to the material. This endeavor

was successful and the biomedical grade of pyrolytic

carbon was rapidly incorporated into the existing heart

valve designs.

Today, pyrolytic carbon (Figures 5 and 6) remains a

popular material available for mechanical heart valves,

being used in more than 4 million implants in more than 25

different valve designs for a clinical experience on the

order of 18 million patient-years.

Pyrolytic carbon (PyC) belongs to the family of turbostratic

carbons, which have a similar structure of graphite, but

subtly different. In graphite, the carbon atoms are

covalently bonded in planar hexagonal arrays that are

stacked and held together by weak interlayer bonding. For

turbostratic carbons, the stacking sequence is disordered,

resulting in wrinkles or distortions within layers. This

structural distortion provides the superior ductility and

durability of pyrolytic carbon, compared to other carbon

structures such as graphite.

Figure 5. Crystal structure of graphite [23].

Mechanical properties [23]

The Pyrolytic carbon, with its inherently dense, glassy

structure and its ability to be highly polished, has become a

popular choice. Furthermore its electrical conductivity is

useful in allowing it to become electrostatically charged so

that it can repel the blood cells. This unique material is one

of the most blood-compatible of all man-made materials, as

opposed to metals. The human body recognizes implanted

metal as a foreign material, and protects itself from the

object by coating it with layers of blood. But, pyrolytic

carbon and other so-called blood-compatible coatings are

unrecognized by the body and are accepted.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 12

Figure 6. Acoustic emission amplitude versus frequency

for crack extensionsg. This plot shows an emission peak at

90kHz, indicating a normal mode crack extension in a

pyrolytic carbon test sample [23].

In its processed form, pyrolytic carbon is a microscopically

smooth, hard, black ceramic-like material. Like ceramic, it

is subjected to brittleness.

Fortunately, pyrolytic carbon possesses a mechanical

property that mitigates this fragility in the presence of

flaws, making it inherently difficult to accidentally

introduce cracks of significant size into the material. In

particular, unlike true ceramics, pyrolytic carbon is highly

ductile. Thus, if a sharp, hard object is pressed into

pyrolytic carbon, it can respond by deforming locally to

accommodate the object elastically. When the object is

withdrawn, there may be no residual depression, and little

or no microcracking surrounding the site. It is this intrinsic,

atomic microstructure-derived resistance to externally

imposed crack nucleation that permits such an otherwise

brittle material to be used in the human body.

The mechanical properties of pyrolytic carbon are largely

dependent on the density as shown in Figures 8 and 9. The

increased mechanical properties are directly related to the

increased density, which indicates that the properties

depend mainly on the aggregate structure of the material. Graphite and glassy carbon have lower mechanical strength

than pyrolytic carbon as given in table 1. However, the

average modulus of elasticity is almost the same for all

carbons. The strength and toughness of pyrolytic carbon are

quite high compared to graphite and glassy carbon. This is

due to the smaller number of flaws and unassociated

carbons in the aggregate.

Figure 7. Fracture stress versus density for unalloyed LTI

pyrolytic carbons [26].

Figure 8. Elastic modulus versus density for unalloyed LTI

pyrolytic carbons [25].

Deposition of pyrolytic carbon coatings for heart valves

For heart valves, a silicon-alloyed pyrolytic carbon is used

in the form of a thick coating on a polycrystalline graphite

substrate. Silicon is added to improve mechanical

properties such as stiffness, hardness, and wear resistance,

without significant loss in biocompatibility. Components

are made by co-depositing carbon and silicon carbide on

the graphite substrate by a chemical vapor-deposition,

fluidized bed process that uses a gaseous mixture of

silicon-containing carrier gas with a hydrocarbon.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 13

Table 1. Properties of various types of carbon [25].

Type

Graphite Glassy Pyrolytic

carbon

Density, g/ml 1.5 -1.9

1.5

1.5 -2.0

Elastic

modulus,

MPa

24

24

28

Toughness,

m-N/ cm3

138

172

517

(525a)

Compressive

strength

6.3

0.6

4.8

a 1.0 w /o Si – alloyed pyrolytic carbon, Pyrolite

(Carbomedics, Austin, Tex).

Pyrolytic Carbon Mechanical Valves in the Market

a. OmnicarbonTM

The Omnicarbon mechanical heart valve is manufactured,

marketed and sold, by Medical CV. Blair Mowery,

president and chief executive officer of Medical CV, noted

that the results from at least 10 clinical studies, including

over 10,000 patient years of use, have consistently

demonstrated one-third to one-half fewer complications

with the Omnicarbon valve, such as blood clots and stroke,

compared to other mechanical valves. The Omnicarbon

heart valve is a monoleaflet valve; a valve with a single

hingeless pivoting disc to employ pyrolytic carbon in both

its housing and disc for improved blood compatibility. Also

Omnicarbon valve does not have fixed pivot recesses that

are characteristic of bileaflet designs and that are

demonstrated to be the primary location for blood clot

formation.

The disc is slightly curved and retained within the housing

ring, located 1800 from each other on the other side of the

housing ring. The disc closes on the housing ring at a 120

angle relative to the plane of the housing ring, and can

open to a maximum angle of 800. The disc rotates freely

within the housing ring because there are no fixed hinges

within the housing ring. Because there are no struts

protruding across the flow orifice, the open disc separates

the flow channel into two orifices.

Indications for Use:

The OmnicarbonTM is indicated for the replacement of

dysfunction, native or prosthetic, aortic or mitral valves.

Contra indications:

The OmnicarbonTM is contraindicated for patients unable to

tolerate anticoagulation therapy.

Figure 8. OmnicarbonTH mechanical heart valve [17].

Adverse Events potentially associated with the use of

mechanical cardiac valves include:

• Angina.

• Cardiac arrhythmia.

• Clinically significant transvalvular regurgitation.

• Disc impingement/ entrapment.

• Endocarditis .

• Heart failure.

• Hemolysis or hemolytic anemia.

• Hemorrhage.

• Myocardial infarction.

• Nonstructural dysfunction.

• Perivalvular leak.

• Stroke.

• Structural dysfunction.

• Thromboembolism.

• Tissue interference with valve function.

• Valve thrombosis.

Precautions and Warnings: In order to avoid harmful

damages to the health of the patient, the following

precautions and warnings must be taken into account:

• Do not use the valve if the use-before-date on the

package has expired.

• If the disc disengages undetected handling

damage or extreme pressure on the disc may

cause this. Should disengagement occur, do not

attempt to re-engage the disc into the valve

housing; the valve should not be implanted.

• If the valve came in contact with blood, do nor

attempt to clean and resterilize such a valve for

use in another person. Foreign protein transfer

and/ or residue from cleaning agents may cause a

tissue reaction.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 14

• Passing a catheter, surgical instrument, or

pacemaker lead through the OmnicarbonTM valve

may cause serious valvular insufficiency, damage

the valve, and/ or cause catheter entrapment.

• Over sizing occurs when too large a valve is

forced into the tissue annulus. This may cause

adjacent tissue to inhibit the free movement and

full travel of the valve disc.

• No hard, sharp instruments should come in

contact with the disc or valve housing ring may

cause scratches or other surface imperfections

which may result in blood injury, thrombus

formation and/ or structural damage.

• A valve soiled by fingerprints or foreign

materials may cause clotting or blood damage.

b. ON-X Carbon [17]

Dr. Vincent Gott compared the clotting tendencies of

silicon carbide, pyrolytic carbon alloyed with silicon

carbide and pure pyrolytic carbon. Pure carbon was shown

to be least thrombogenic. MCRI overcame the need for

silicon carbide by applying new technologies to the

pyrolytic carbon manufacturing processes. Without silicon

carbide, the pure carbon’s surface finish is unmatched in

purity and smoothness. There was an additional reward in

purifying carbon. ON-X carbon is 50% stronger than

previous carbons. Its added flexural strength is essential to

the manufacturability of the ON-X valve’s sophisticated

design.

How does ON-X work?

Like natural valves, mechanical heart valves are one-way

valves that are opened and closed by the action of the blood

pushing on flaps known as leaflets. The ON-X valve’s

leaflets (Figure 9) are somewhat like double doors that

open and close but never latch. In the case of doors, if the

wind blows from one direction, the doors will be blown

open. If blown from the other direction, the doors will be

close. This analogy is an over simplification as the

demands of the body can be both rigorous and subtle.

Safety and efficiency of ON-X

Two measures of a good mechanical heart valve are safety

and efficiency. To be safe, a valve must not wear out, break

or malfunction. It must not be ejected by the body’s

immune system. It must minimize blood clotting and

damage to the blood with minimum amount of effort.

ON-X and common safety problems

Thrombosis is a problem that causes blood clots on the

working surface of a valve, which impairs valve function.

As a thrombus gets bigger, it will eventually block the

moving parts so that the valve can no longer open and / or

close fully. By using extremely smooth ON-X carbon and

by designing the valve to induce smooth flow and thorough

self-cleaning, ON-X reduces the risk of thrombosis.

In the case of tissue encroachment, the body’s healing

process can also impair valve function. As the body heals

around the mechanical heart valve, tissue builds up around

the valve. This becomes a problem if the tissue grows over

the valve and begins to block it or restrict the moving parts.

The ON-X valve was designed with leaflet guards and

optimized length to ensure that tissue doesn’t interfere with

valve function.

In the blood damage problem, turbulence and rapid

changes in pressure can affect the blood flow. The longer

flared body of the ON-X smoothes flow. The leaflets are

free to open completely to align with the flow and only

move a short distance to close, which reduces turbulence

and buffers. The ON-X valve minimizes damage to blood

cells.

Typical mechanical properties of ON-X Carbon are given

in table 2.

Figure 9. ON-X mechanical heart valve [17].

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 15

Table 2. Typical surface and mechanical properties of On-

X Carbon [9].

66.. TTiittaanniiuumm ((TTii)) [[2244 aanndd 2277]]

TThhee hhiigghh ssttrreennggtthh,, llooww wweeiigghhtt,, oouuttssttaannddiinngg ccoorrrroossiioonn

rreessiissttaannccee ppoosssseesssseedd bbyy ttiittaanniiuumm aanndd ttiittaanniiuumm aallllooyyss hhaavvee

lleedd ttoo aa wwiiddee aanndd ddiivveerrssiiffiieedd rraannggee ooff ssuucccceessssffuull

aapppplliiccaattiioonnss wwhhiicchh ddeemmaanndd hhiigghh lleevveellss ooff rreelliiaabbllee

ppeerrffoorrmmaannccee iinn ssuurrggeerryy aanndd mmeeddiicciinnee.. MMoorree tthhaann 11000000

ttoonneess ((22..22 mmiilllliioonn ppoouunnddss)) ooff ttiittaanniiuumm ddeevviicceess ooff eevveerryy

ddeessccrriippttiioonn aanndd ffuunnccttiioonn aarree iimmppllaanntteedd iinn ppaattiieennttss

wwoorrllddwwiiddee eevveerryy yyeeaarr..

Table 3. Biocompatibility Tests & Results [9].

TThhee ttiittaanniiuumm aallllooyy TTii--66AAll--44VV iiss uusseedd aass tthhee ccaarrrriieerr

ssttrruuccttuurree ffoorr rreeppllaacceemmeenntt hheeaarrtt vvaallvveess.. TThhee ttiittaanniiuumm iiss rriinngg

sshhaappeedd aanndd ssuuppppoorrttss tthhee mmoovviinngg mmeecchhaanniissmmss ooff tthhee

rreeppllaacceemmeenntt vvaallvvee.. IItt aallssoo ccaarrrriieess tthhee ppoollyyeesstteerr ssttrruuccttuurree

tthhaatt bbiinnddss tthhee vvaallvvee ttoo tthhee ttiissssuuee..

MMeeddiiccaall ggrraaddee ttiittaanniiuumm aallllooyyss hhaavvee aa ssiiggnniiffiiccaannttllyy hhiigghheerr

ssttrreennggtthh ttoo wweeiigghhtt rraattiioo tthhaann ccoommppeettiinngg ssttaaiinnlleessss sstteeeellss..

TThhee rraannggee ooff aavvaaiillaabbllee ttiittaanniiuumm aallllooyyss eennaabblleess mmeeddiiccaall

ssppeecciiaalliissttss’’ ddeessiiggnneerrss ttoo sseelleecctt mmaatteerriiaallss aanndd ffoorrmmss cclloosseellyy

ttaaiilloorreedd ttoo tthhee nneeeeddss ooff tthhee aapppplliiccaattiioonn.. TThhee ffuullll rraannggee ooff

aallllooyyss rreeaacchheess ffrroomm hhiigghh dduuccttiilliittyy ccoommmmeerrcciiaallllyy ppuurree

ttiittaanniiuumm uusseedd wwhheerree eexxttrreemmee ffoorrmmaabbiilliittyy iiss eesssseennttiiaall,, ttoo

ffuullllyy hheeaatt ttrreeaattaabbllee aallllooyyss wwiitthh ssttrreennggtthh aabboovvee 11330000 MMPPaa

((119900 kkssii)).. SShhaappee––mmeemmoorryy aallllooyyss bbaasseedd oonn ttiittaanniiuumm ffuurrtthheerr

eexxtteenndd tthhee rraannggee ooff uusseeffuull pprrooppeerrttiieess aanndd aapppplliiccaattiioonnss.. AA

ccoommbbiinnaattiioonn ooff ffoorrggiinngg oorr ccaassttiinngg,, mmaacchhiinniinngg aanndd

ffaabbrriiccaattiioonn aarree tthhee pprroocceessss rroouutteess uusseedd ffoorr mmeeddiiccaall

pprroodduuccttss..

Functional Requirements

The following requirements must be satisfy to use titanium

as a biomaterial in heart valves implants:

• The body’s immune system must not attack the

biomaterial.

• Compatible with body tissues and fluids

• Must has strength, flexibility and hardness

• Must be nontoxic, nonreactive or biodegradable

• The valve must also be anchored to the inside of

the heart.

• The material must not exhibit mechanical fatigue

over the device’s lifetime.

• The material’s surface must have an acceptable

low propensity for thrombus formation, as well as

the best possible blood compatibility

• Must not be prone to calcification.

Tests Results

Cytotoxicity L-929

Membrane Elution

non-cytotoxic

Sensitization ISO

Kligman

0% sensitization:

Grade I

sensitization

rate, not significant

Irritation Saline CSO negligible irritant

Acute Systemic

Toxicity Saline CSO

negative

Rabbit Pyrogen non-pyrogenic

USP Physical / Chemical

Screening Tests

passes USP

standards

Mutagenicity Ames non-mutagenic

Hemolysis Direct Contact

Rabbit Blood

non-hemolytic

Complement Activation non-activating

Property Units On-X

Wear Resistance

mm3/km, 10-6 <1.23

Coefficient of Friction

------ 0.15

Young's Modulus

GPa 26

Flexural Strength

MPa 490

Density

gm/cm3 1.9

Strain to Failure

% 1.6

Strain Energy

MPa-mm/mm 7.7

Residual Stress

MPa 18.2

Fracture Toughness

MPa m1/2 1.67

Fatigue Threshold

m/cycle

(DK70.3)

1.11

Fatigue Crack Velocity m/cycle, 10-15

3.98

Critical Surface Tension

dynes(cm) 42

Surface Roughness

Ra(nm) 33.9

Surface Chemistry

Carbon

Atomic % ~85

Surface Chemistry

Silicon

Atomic % 0

Surface Chemistry

Oxygen

Atomic % ~15

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 16

Titanium Performance in Medical Applications

The titanium alloy Ti-6Al-4V is classified as biologically

inert biomaterial or bioinert. TTiittaanniiuumm iiss jjuuddggeedd ttoo bbee

ccoommpplleetteellyy iinneerrtt aanndd iimmmmuunnee ttoo ccoorrrroossiioonn bbyy aallll bbooddyy

fflluuiiddss aanndd ttiissssuuee,, aanndd iiss tthhuuss wwhhoollllyy bbiiooccoommppaattiibbllee.. As

such, it remains essentially unchanged when implanted into

human bodies because of its excellent corrosion resistance.

The human body is able to recognize bioinert materials as

foreign, and tries to isolate them by encasing them in

fibrous tissues. However, they do not illicit any adverse

reactions and are tolerated well by the human body.

Furthermore, they do not induce allergic reactions such as

has been observed on occasion with some stainless steels,

which have induced nickel hypersensitivity in surrounding

tissues.

TThhee ffaavvoorraabbllee cchhaarraacctteerriissttiiccss ooff ttiittaanniiuumm iinncclluuddiinngg

iimmmmuunniittyy ttoo ccoorrrroossiioonn,, bbiiooccoommppaattiibbiilliittyy,, ssttrreennggtthh,, llooww

mmoodduulluuss aanndd ddeennssiittyy.. TThhee lloowweerr mmoodduulluuss ooff ttiittaanniiuumm aallllooyyss

ccoommppaarreedd ttoo sstteeeell iiss aa ppoossiittiivvee ffaaccttoorr.. TTwwoo uusseeffuullnneessss

ppaarraammeetteerrss ooff tthhee iimmppllaannttaabbllee aallllooyy aarree tthhee nnoottcchh

sseennssiittiivviittyy.. TThhee rraattiioo ooff tteennssiillee ssttrreennggtthh iinn tthhee nnoottcchheedd

vveerrssuuss uunn--nnoottcchheedd ccoonnddiittiioonn aanndd tthhee rreessiissttaannccee ttoo ccrraacckk

pprrooppaaggaattiioonn,, oorr ffrraaccttuurree ttoouugghhnneessss.. TTiittaanniiuumm ssccoorreess wweellll iinn

bbootthh ccaasseess.. TTyyppiiccaall NNSS//TTSS rraattiiooss ffoorr ttiittaanniiuumm aanndd iittss aallllooyyss

aarree 11..44 -- 11..77 ((11..11 iiss aa mmiinniimmuumm ffoorr aann aacccceeppttaabbllee iimmppllaanntt

mmaatteerriiaall)).. FFrraaccttuurree ttoouugghhnneessss ooff aallll hhiigghh ssttrreennggtthh

iimmppllaannttaabbllee aallllooyyss iiss aabboovvee 5500 MMPPaamm--11//22

wwiitthh ccrriittiiccaall ccrraacckk

lleennggtthhss wweellll aabboovvee tthhee mmiinniimmuumm ffoorr ddeetteeccttiioonn bbyy ssttaannddaarrdd

mmeetthhooddss ooff nnoonn--ddeessttrruuccttiivvee tteessttiinngg.. The two most common

types of Ti-6Al-4V used for the implants are Ti-6Al-4V

Grade 5 and Grade 23.

Ti-6Al-4V (Grade 5)

This alpha-beta alloy is the workhorse alloy of the titanium

industry. The alloy is fully heat treatable in section sizes up

to 15 mm and is used up to approximately 400°C (750°F).

Since it is the most commonly used alloy – over 70% of all

alloy grades melted are a sub-grade of Ti6Al4V.

The addition of 0.05% palladium (grade 24), 0.1%

ruthenium (grade 29) and 0.5% nickel (grade 25)

significantly increases corrosion resistance in reducing

acid, chloride and sour environments, raising the threshold

temperature for attack to well over 200°C (392°F).

Ti-6Al-4V (Grade 23)

The essential difference between Ti6Al4V ELI (grade 23)

and Ti6Al4V (grade 5) is the reduction of oxygen content

to 0.13% (maximum) in grade 23. This offers improved

ductility and fracture toughness, with some reduction in

strength. Grade 29 also having lowered level of oxygen will

deliver similar levels of mechanical properties to grade 23

according to processing.

Advancements: Titanium

Material selection for implantable medical devices has

improved with the availability of nitinol, or “NiTi”, a

nickel-titanium alloy that has proved to be biocompatible,

durable and non-thrombogenic.

Researchers at University of California, Los Angeles

(UCLA) have designed thin-film NiTi semi-lunar heart

valve for use in both surgical and non-surgical

(transcatheter) human heart valve replacements:

• Surgically implantable thin-film NiTi valves are

undergoing in vitro testing to determine their

functionality, durability and corrosive properties;

• Designs and prototypes of percutaneously

inserted catheter-based thin-film NiTi valves are

under continuing development.

• Use of thin-film nitinol as a novel material for the

development of improved human prosthetic heart

valves for surgical implantation and for

percutaneous insertion.

7. Biomaterials versus stainless steel (Table 4)

Alumina is a versatile material with applications in

medicine, because it has good compatibility with human

environment. Compared to stainless steel, mechanical

properties of Alumina are worst in modulus of elasticity,

shear modulus, thermal expansion coefficient. Alumina is

better in stress, strain and safety factor. Stainless steel is

more resistant but it is not utilized on heart valve implants,

because it doen not have good biocompatibility with human

blood. Alumina is a bioceramic while stainless steel is a

biometal with different density.

Stainless steel is stiffer than the titanium alloy Ti6Al4V.

This is explained by the steel’s higher modulus of elasticity

(196 GPa vs. 120 GPa). Steel is more rigid than titanium

(steel’s higher shear modulus: 80 GPa vs. 44 GPa).

Stainless steel is more fracture resistant than titanium (steel

has more tensile strength: 875 MPa vs. 616 MPa). Titanium

is more resistant to yielding that stainless steel (Ti has a

higher yield stress: 950 MPa vs. 700 MPa). See appendix

VI.

Pyrolytic carbon is less stiffer than stainless steel. This is

explained by the steel’s higher modulus of elasticity (196

GPa vs. 17-28 GPa). Stainless steel is more resistant than

pyrolytic carbon because of its tensile strength is biggest

compared to pyrolytic carbon (875 MPa vs. 200 MPa) .

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 17

Stainless steel is stiffer than the polyetherurethane. This is

explained by the steel’s higher modulus of elasticity (196

GPa vs. 0.016 GPa). Stainless steel is more fracture

resistant than polyetherurethane (steel has more tensile

strength: 875 MPa vs.49.7 MPa).

Stainless steel is stiffer than Polyester. This is explained by

the steel’s higher modulus of elasticity (196 GPa vs. 1.84

GPa). Polyester is less fracture resistant than Stainless steel

(stainless steel has more tensile strength: 875MPa vs.

48.3MPa). Stainless steel is more resistant to yielding that

polyester (Stainless steel has a higher yield stress: 700 MPa

vs. 59.3MPa).

SUMMARY

When designing prosthetic heart valves, there are several

characteristics of natural heart valves one aims to mimic.

These include minimal transvalvular pressure gradients,

minimal regurgitation fractions, central flow characteristics

and complete biocompatibility. The materials for the

prosthesis should be durable, non-toxic and

nonthrombogenic; ideally, the materials should not require

the long-term use of anticoagulant therapy. The prosthetic

heart valve should be surgically implanted with ease and

not interfere with normal cardiac function and anatomy.

The normal function of the prosthetic valve should be

quiet, should not damage cellular blood elements or cause

denaturing of proteins. Finally, prosthetic valves should be

readily available, manufactured with ease and relatively

inexpensive.

Cardiovascular surgeons must weigh the advantage of the

durability of mechanical type prostheses without the need

for long-term anticoagulant therapy. Therefore, physicians,

biomedical engineers and other inventors have yet to

design the “ideal” prosthetic valve substitute.

One approach to meet the characteristics of the “ideal”

prosthetic valve includes new fixing processes for

bioprosthetic valves that greatly improve durability,

decrease the incidence of dystrophic calcification and do

not change the relatively nonthrombogenic nature of

existing bioprostheses. Another approach would be the

introduction of a new durable, nonthrombogenic material

Property Units Alumina Titanium Polyetheruretane Pyrolitic

Carbon

Stainless

Steel

Polyester

Poisson’s

Ratio

N/A 0.33 0.33 0.40 0.3-0.4 0.27-0.30 0.33-0.49

Hardness GPa 20.6 2.24 50 10 5-8.5 90

Young’s

Modulus

GPa 392 120 0.16 17-28 196 1.84

Shear

Modulus

GPa 163 44 1-2 N/A 75-80 0.744-1.586

Tensile

Strength

MPa 637 616 49.7 200 875 48.3

Compressive

Stress

GPa 4900 N/A 50-70 900 N/A 59.6

Yield Stress MPa 15.4*103 950 11.9 100 700 59.3

Ultimate

Stress

MPa 119 930 50 N/A 59.3 46.5

Coefficient

of Thermal

Expansion

(Linear)

10-6

per °C

6.2 11.9 25 10 N/A 70

Table 4. Mechanical Properties of biomaterials

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 18

from which mechanical type prosthesis could be fashioned.

Some think a synthetic valve which could be constructed

to copy the durability and higher than acceptable

transvalvular pressure gradients.

Currently available prostheses are markedly improved over

some earlier valve substitutes, but the search for the “ideal”

prosthetic valve continues. To advance towards this goal, a

major breakthrough in either materials science or collagen

biochemistry must occur.

ACKNOWLEDGEMENTS

We would like to thank Dr. Megh R. Goyal, University of

Puerto Rico, for his advice. The authors also to thank Yi-

Ren Woo (St. Jude Medical Principal engineer) and Carlos

Rosario, M.D. (Mayaguez Cardiology), Prof. Pablo

Cáceres, at University of Puerto Rico for providing

technical information.

REFERENCES

1. http://cape.uwaterloo.ca/che100projects/heart/files/test

ing.htm

2. http://titaniuminfogroup.co.uk

3. http://www.azom.com

4. http://www.azom.com/details.asp?ArticleID=105

5. http://www.azom.com/details.asp?ArticleID=2103

6. http://www.ceramics.nist.gov/srd/summary/scdaos.htm

7. http://www.cnn.com/HEALTH/library/DS/00421.html

8. http://www.domme.ntu.ac.uk/research/biomec/pap...ri

alchoice.html

9. http://www.fda.com

10. http://www.gla.ac.uk/departments/cardiacsurgery/biog

_Berraca.htm

11. http://www.icr-heart.com/journal/unalloyed_pyrolytic

_carbon_for_i.htm

12. http://www.lib.umich.edu/dentlib/Dental_tables/toc.html

13. http://www.library.drexel.edu/research/guides/pdfs/ma

terialproperties.html

14. http://www.library.drexel.edu/research/guides/pdfs/ma

terialproperties.html#bio

15. http://www.mcritx.com/carbon_properties.htm

16. http://www.medhelp.org/forums/cardio/archive/848.html

17. http://www.medtronics.com

18. http://www.mkt-intl.com/ceramics/aluminaphotos.htm

19. http://www.mst.dk/udgiv/Publications/1999/87-7909-

416-3/html/bil03_eng.htm

20. http://www.polymertech.com/materials/elasthane.html

21. http://www.pyrocarbon.com

22. http://www.sts.org

23. http://www.swri.org/3pubs/ttoday/summer99/valve.htm

24. http://www.titanium.org

25. Park, J.B. 1984. Biomaterials Science and

Engineering. pp.212-216, 252-256 New York: Plenum

Press, New York.

26. Mitamura Y, Hosooka K, Matsomoto T, Otaki K and

Sakai K. Development of a fine ceramic Heart valves.

Journal of Biomaterials Application. Publisher Sage

Publication, London.

27. Sharma, Szycher. 1991. Blood Compatible Materials

and Devices. Technomic Publishing Company. Inc.

pp.33, 156-163.

GLOSSARY

Alloy: a material that consisting of two or more metals or a

metal and non-metal.

Anesthesiologists: a physician specializing in

anesthesiology (anesthesia is used during some procedures

and surgery).

Anticoagulant: a drug used to thin the blood, keeps blood

from clotting.

Aorta: the largest artery in the human body, it carries

blood from the heart to every part of the body.

Aortic valve: the valve between the left ventricle and the

aorta.

Atria (Atrium): upper-receiving chambers of the heart

Calcification: formation of calcium deposits on the surface

of the material.

Cardiac catheterization: a highly specialized non-surgical

technique that allows cardiologists to examine coronary

arteries for blockage using thin catheters inserted into the

heart.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 19

Cardiologists: a physician specializing in the heart.

Cardio thoracic Surgeons: heart surgeons.

Creep: additional strains develop, when loaded for long

period of time is applied.

Diastole: during the cardiac cycle, when the heart relaxes

& allows blood to flow in.

Ductile: capable of being drawn out into a thin wire or

thread.

Hingeless : do not have a movable joint by means of which

it can turn on the frame.

Infarct: dead tissue as a result of obstructed blood flow to

the area.

Minimally invasive surgery: techniques that use small

incisions to gain access to the surgical site.

Mitral valve: the valve that separates the left atrium from

the left ventricle.

Modulus of Elasticity: slope of the straight line from the

stress- strain diagram.

Pericardial valve: tissue valve made from bovine tissue.

Poisson’s Ratio: ratio of the lateral and axial strain,

property of materials.

Polish: to make or become smooth and glossy by rubbing.

Porcine valve: tissue valve made from a pig’s aortic heart

valve.

Pulmonary valve: the valve between the right ventricle

and the pulmonary artery.

Regurgitation: backward flow of blood due to inability of

valve to work properly.

Restenosis: recurrence of the blockage or narrowing of the

artery or valve.

Stenosis: narrowing of artery or heart valves.

Strain: elongation per unit of length.

Stress: intensity of force per unit of area.

Systole: in the cardiac cycle, when the heart contracts

(pumps).

Thrombus: a blood clot.

Valves: flap-like structures, which control the flow of

blood through the heart.

Valve surgery: there are a total of 4 valves in the human

heart, valve surgery repairs or replaces damaged or scarred

valves.

Ventricle: the large lower pumping chamber of the heart.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez 20

APPENDIX I: FDA POLICIES OF MECHANICAL

MITRAL VALVE REPLACEMENT. [9]

1. ATS Open Pivot® Bileaflet Heart Valve

This is a brief overview of information related to FDA's

approval to market this product. See the links below to the

Summary of Safety and Effectiveness and product labeling

for more complete information on this product, its

indications for use, and the basis for FDA's approval.

Product Name: ATS Open Pivot® Bileaflet Heart Valve

Manufacturer: ATS Medical, Inc.

Address: 3905 Annapolis Lane, Suite 105, Minneapolis,

Minnesota 55447

Approval Date: October 13, 2000

Approval Letter:

http://www.fda.gov/cdrh/pdf/p990046a.pdf

What is it? The ATS Open Pivot® Bileaflet Heart Valve is

a mechanical heart valve with two leaflets (flap like

structures) in the shape of a circle, each leaflet a half the

circle, surrounded by a ring made of polyester fabric. The

leaflets are made of carbon. The valve is used to replace a

patient’s own aortic or mitral valve, or another prosthetic

aortic or mitral valve.

How does it work? The ATS Open Pivot® Bileaflet Heart

Valve uses two half discs (bileaflets) that open and close as

blood flows through the valve to operate like the patient’s

natural heart valve. (Heart valves control the blood flow

through the chambers of the heart.)

When is it used? The ATS Open Pivot® Bileaflet Heart

Valve is intended to replace diseased, damaged, or

malfunctioning natural or prosthetic aortic or mitral valves.

Heart valves may not always work as well as they should.

Disease or other heart valve malfunction may cause the

heart valve tissue to thicken, harden, weaken, or stretch. If

the valve fails to open and close properly, it can block or

interfere with blood flow causing a decrease in the efficient

flow of blood through the heart. This can reduce a patient’s

quality of life.

What will it accomplish? A patient who has a diseased,

damaged, or malfunctioning heart valve may feel weak,

tired, or otherwise handicapped. Surgical replacement of

the affected heart valve may be an effective option to

improve the patient’s quality of life.

When should it not be used? The valve should not be used

in patients who are unable to tolerate anticoagulant therapy

or the use of blood-thinning drugs.

2. Subpart D--Cardiovascular Prosthetic Devices

Sec. 870.3925 Replacement heart valve.

(a) Identification. A replacement heart valve is a device

intended

to perform the function of any of the heart`s natural valves.

This device includes valves constructed of prosthetic

materials, biologic valves (e.g., porcine valves), or valves

constructed of a combination of prosthetic and biologic

materials.

(b) Classification. Class III (premarket approval).

(c) Date premarket approval application (PMA) or notice of

completion of a product development protocol (PDP) is

required. A PMA or

a notice of completion of a PDP is required to be filed with

the Food and Drug Administration on or before December

9, 1987 for any replacement heart valve that was in

commercial distribution before May

28, 1976, or that has on or before December 9, 1987 been

found to be substantially equivalent to a replacement heart

valve that was in commercial distribution before May 28,

1976. Any other replacement heart valve shall have an

approved PMA or a declared completed PDP in effect

before being placed in commercial distribution.

[45 FR 7907-7971, Feb. 5, 1980, as amended at 52 FR

18163, May 13, 1987; 52 FR 23137, June 17, 1987]

3. Subpart D--Cardiovascular Prosthetic Devices

Sec. 870.3945 Prosthetic heart valve sizer.

(a) Identification. A prosthetic heart valve sizer is a device

used to measure the size of the natural valve

opening todetermine the size of the appropriate

replacement heart valve.

(b) Classification. Class I (general controls). The device is

exempt from the premarket notification procedures in

subpart E of part 807

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 21

APPENDIX II: OTHER PATENTS FOR COMPANY'S PRODUCTION.

APPLICATION

NUMBER / DATE of

APPROVAL

DEVICE TRADE NAME COMPANY NAME

CITY, STATE, &

ZIP

DEVICE DESCRIPTION /

INDICATIONS

P810002/S056

5/3/01

Real-Time

St. Jude Medical® Mechanical

Heart Valve – Mater Series

Coated Aortic Valved Graft,

Model CAVGJ-514 00

St. Jude Medical, Inc.

St. Paul, MN

55117

Approval for the changes to the country

of origin of the bovine collagen used

for the Hemashield® graft, which is a

component of the device. The device,

as modified, will be marketed under the

trade name St. Jude Medical®

Mechanical Heart Valve – Master Series

Coated Aortic Valved Graft, Model

CAVGJ-514 00 in sizes 19, 21, 23, 25,

27, 29, 31, 33 mm, and is indicated for

the replacement of the aortic valve and

the ascending aorta.

P000037

5/30/01

On-X® Prosthetic Heart

Valve, Model ONXA

Medical Carbon Research

Institute, LLC

Austin, TX

78754

Approval for the On-X® Prosthetic

Heart Valve, Model ONXA in the aortic

position including sizes 19, 21, 23, 25,

and 27/29 mm. This device is indicated

for replacement of diseased, damaged,

or malfunctioning native or prosthetic

heart valves in the aortic position.

P790018/S031

11/24/97

Medtronic Hall™ Prosthetic

Heart Valve (Models A7700

and M7700)

Medtronic Heart Valves,

Inc.

Irvine, CA

92714

Approval for a modification to the

controlled environment area for certain

manufacturing steps.

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 22

APPENDIX III: POLYURETHANE PROPERTIES.

Trade name

CARDIOMAT

610

MITRATHANE

2007

PAMPUL-3-

AMEO

Pur 1025/1

Producer

KONTRON

Cardiovascular

Inc.

MITRAL

MEDICAL

Internacional Inc.

BEIERSDORF

AG

ENKA AG

Pur - type

Polyetherurethane

Segm.

Polyetherurethane

urea

Polyetherurethane

Polyesterurethane

CONCENTRATION

(%)

15

25

10

15

SOLVENTS

THF/ DIOXAN:

2/1

DMAC

DMAC

DMAC

VISCOSITY

( m Pa s) at °C

2010 at 30°C

77500 at 23°C

4200 at 23°C

48500 at 23°C

DENSITY ( g/cm^3 )

1.11 + 0.03

?

?

1.12

HARDNESS

( Shore A)

80

65 + 5

75 + 5

87

TENSILE

STRENGTH

( N/mm^2)

28.0

39.2 + 5.0

49.0 + 7.0

50.6

ELONGATION at

BREAK ( %)

500

775 + 50

605 + 30

649

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 23

APPENDIX IV: ALUMINA MITRAL VALVE AND THROMBUS FORMATION COMPARED WITH OTHER

MATERIALS.

ALUMINA MITRAL VALVE

THROMBUS FORMATION

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 24

APPENDIX V: ADVERSE EVENTS OF DIFFERENT HEART VALVE IMPLANTATION.

Early Postoperative Adverse Events

n (% or cases)

Event

AVR (125)

MVR (70)

DVR (37)

Death, all causes

Thromboembolism, AII

Thromboembolism, TIA

Thromboembolism, Nontransient

Valve Thrombosis

Anticoagulant – Related Hemorrhage, major

Endocarditis

Perivalvular Leak, major

Pannus Tissue Interference

Hemolytic Anemia

Structural Failure

Unacceptable Hemodynamics

Other Nonstructural Dysfuction

Reoperation

Explantation

4 (3.2)

5 (4.0)

2 (1.6)

3 (2.4)

0

0

1 (0.8)

1 (0.8)

0

0

0

0

0

2 (1.6)

2 (1.6)

6 (8.6)

0

0

0

1 (1.4)

1 (1.4)

0

0

0

0

0

0

0

1 (1.4)

1 (1.4)

2 (5.4)

0

0

0

0

0

0

1( 2.7)

0

0

0

0

0

1 (2.7)

0

Abbreviations; n= number of patients

AVR = aortic valve replacement, MVR = mitral valve replacement, DVR = double valve replacement

TIA = transient ischemic attack

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 25

APPENDIX VI: MECHANICAL PROPERTIES COMPARISON WITH OTHER BIOMATERIALS.

APPENDIX VII: TABLE OF TITANIUM YIELD STRENGTH VERSUS DENSITY.

Yield Strength vs Density

0 20 40 60 80 100 120 140 160 180 200

Yie

ld S

tren

gth

(M

Pa)

Density (g/cm^3)

Ti6Al4V

Ta

316L SS

CoNiCrMo

F 562

Annealed

Annealed

Annealed

Cold-worked

Cold-worked

As-cast

Wrought annealed

Cold-worked

Unalloyed Grade 4

Heat treated

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 26

APPENDIX VIII: EXERCISES

1. AXIALLY LOADED MEMBERS: A polyester graft supports a tensile load of 48 lb when the heart beats. The inner

and outer diameter of graft are d1 = 0.143 in. and d2 = 0.169 in. respectively, his length is 0.5 in. The elongation due

to the load is 0.018 in. Find the stress and strain.

Solution:

Given:

d1 = 0.143 in

d2 = 0.619 in

L = 0.5 in

P = 48 lb

Calculate the cross-sectional area:

A = ( π / 4 ) ( d2² - d1² ) = ( π / 4 ) ( 0.169² - 0.143² ) = 0.00637 in. ²

Calculate the stress:

σ = P / A = 48.0 lb / 0.00637 in. ² = 7,535 psi

Calculate the strain :

∈ = δ / L = 0.018 in. / 0.5 in. = 0.036

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 27

2. AXIALLY LOADED MEMBERS: A stainless steel implant of 2.0 in is put inside a wrist bone. Its temperature rises of 27ºC

to 45ºC. Calculate the thermal strain and temperature-displacement relation of the metal. Assume the thermal expansion

coefficient (α) of the implant is 0.000007/ºC.

Bone

Solution:

Given:

T1 = 27ºC

T2 = 45ºC

L = 2.0 in

α = 0.000007/ºC

Є T = α(T2 - T1)

= (0.000007/ºC)(45ºC - 27ºC)

= (0.000007/ºC)(18ºC)

= 1.26 x 10-4

δ T = L Є T

= (2.0in)(1.26 x 10^-4) = 2.52 x 10-4 inch

metal

implant

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 28

3. TORSION: A metallic bar made of stainless steel is implanted in the human vertebral column. The bar has a

diameter d=0.025 m, length L=0.3075m and shear modulus of elasticity. The bar is subjected to torque T, acting at

the ends.

a. If it has a load of 1,000N at .15375 m, calculate the magnitude of the torque.

b. Using the results of the part a, what is the maximum shear stress in the bar?

c. What is the angle of twist between the ends?

d. If the allowable shear stress is 2.037 MPa and the allowable angle of twist is 2.5o rad, what is the maximum

permissible torque?

L= 0.3075m

Solution:

Given:

d = 0.025 m

L = 0.3075 m

G = 196*106 Pa

P = 1000 N allow = 2.037 MPaح

�allow = 2.5o rad

a. T= Pd

= ( 1000N)(0.025m) = 25 Nm

b. حmax = 16T/ πd3

= (16)(25N.m)/ π(0.025m)3 = 8.1487 MPa

c. Ip= πd4/32

= π (0.025m)4 /32 = 38.3 nm

4

� = TL/ GIp

= (25Nm)(0.3075m)/(196x10^6 Pa)(38.3x10^-9m4) = 1.2

º rad

d. T1= πd3 allow/16ح

allow= P/A =1,000N /[(π/4) (0.025m)ح 2] = 2.037 MPa

T1= π (0.025m)3(2.037x10^6 Pa)/16 = 6.25 Nm

T2= GIP�allow/L

= [(196x10^9N/m2)(38.4x10^-9m

4(2.5º)( π rad/180º)] / 0.3075m = 1068 Nm

The maximum permissible torque is smaller of T1 and T2.

T1=6.25 Nm

d = 0.025m

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 29

4. COMPRESSION: An open alumina mitral valve has an inner radius of 0.0508 m and outer radius of 0.0635 m. It is loaded by

a compression blood flow equal to 1,750 mm Hg. A) Determine the force to the valve sectional area. B) Find the maximum shear

stress on this mitral valve prosthesis. ***Hint: Use correct units.

P = 1,750mmHg

Free Body Diagram

Solution:

Presion = Force/Area Force = Area x Presion

Force = (1,750 mmHg) π(.0635+. 0508)2

Force =22.86 mm Hg-m2

22.86 mm Hg (101.325KPa/ 760 mmHg) = 3.048 KPa

Maximum Shear Stress = (V*Q)/(I*b) = (4*P)/(3*π) =

[4 (3.048 KPa)/ 3π] * [(R22+R2R1+R12)/(R24-R14)]

12.192 [(0.06352 + 0.1143 + 0.05082)/ (0.06354 - 0.05084)] =

Maximum stress = 1.47/9.6x10-6= 153.135 KPa

December 2003 Applications of Engineering Mechanics in Medicine, GED at University of Puerto Rico, Mayagüez. 30

5. TENSION, COMPRESSION AND SHEAR: An artificial mitral valve has a carrier structure made of a titanium Ti6Al4V ring.

The ring has a diameter of 30 mm. This ring is formed into a round bar to be used in the lab for experiments with a tensile test

machine. The bar is stretched to a final length of 190 mm. Find the tensile load applied to the specimen and the dilatation.

Ti6Al4V Poisson’s Ratio = 0.33

Ti6Al4V Young Modulus = 120 GPa

Ti6Al4V Yield Stress = 950 MPa

Titanium ring:

1.5 mm

Solution:

a. Initial Length of Bar:

Circumference =2*pi*(30)

= 188.50 mm

b. Strain:

(Lf – Li) / Li = (190-188.5)/188.5

= 0.0079

c. Normal Stress= E*Strain

= (120*10^9)*(0.0079)

= 948*10^6 Pa

d. Hook’s valid?

948*10^6 Pa is less than Yield Stress (950*10^6 Pa)

e. Tensile Load (P) = normal stress*Area

P = 948*10^6 Pa*(pi/4)*(0.0015)^2 = 1675.25 N (Tension)

f. Dilatation (e) = Strain/(1-2V)

= 0.0079*(1-2(0.33))

= 0.002686*100

= 0.2686 %