Kutz M. Handbook of materials selection

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2 ORTHOPEDIC BIOMATERIALS: TOTAL HIP ARTHROPLASTY 1179

Table 4 Summary of Mechanical Properties for Materials Commonly Used in Current

Total Joint Replacement Designs

Property HMWPE Titanium Alloy Co–Cr Alloy Alumina Cortical Bone

Elastic Modulus

[GPa]

2.2 110 220–234 350–400 10– 20

Comp. Strength

[MPa]

4000 130–280

Tensile Strength

[MPa]

3 860 600–1000 270 80–160

Endurance Limit

[MPa] at 10

cycles

620 500

Density [g/ cm

] 0.93–0.94 4.5 9.2 3.9 1.8

Hardness [MPa] 3500 3000–4000 20,000

Titanium alloy, Ti6Al4V; Co–Cr alloy, wrought CoNiCrMo.

Source: Data from ASTM-F136 (1998), ASTM-F562 (2000), ASTM-F648 (2000), Bernache-

Assolant (1991), Brunski (1996), Cowin (1989), and Park and Lakes (1992)

mechanical properties of these common materials and lists cortical bone for

comparison purposes.

In the selection of metals, cobalt–chromium alloys with molybdenum have

been preferentially chosen over those with nickel. Despite the fact that the me-

chanical properties are slightly reduced in the molybdenum alloy compared to

the nickel variant, it has been shown to have improved wear properties. Titanium

is often chosen due to the fact that its elastic modulus is half of that seen in

cobalt–chromium or stainless steel alloys, therefore reducing the stiffness of the

implant and the accompanying bone loss due to stress shielding.

During the past decade, substantial research has been conducted to develop

different pairs of materials for the bearing surface in order to reduce the wear

phenomenon commonly seen in HMWPE, which often leads to premature failure

of an implant. Metal-on-metal and ceramic-on-ceramic head–cup combinations

have shown promise in terms of long-term outcomes, especially when used in

younger patients for whom an implant life of greater than 20 years would be

desirable (Delaunay, 2000; Skinner, 1999; Wagner and Wagner, 2000). The ce-

ramics typically selected have been alumina and zirconia, while the metal is

generally a cobalt–chromium alloy. Titanium does not lend itself to metal–metal

bearing surfaces due to the fact that it tends to seize when in contact with other

metals. The success of these material combinations is due in large part to the

improvements in materials processing, machining, and polishing that have been

achieved in the last quarter century, allowing for the development of high-quality

materials, excellent geometric matches, and highly polished surfaces. This pro-

gress also indicates that designs that were tried out at an earlier point in history,

and were perhaps abandoned, may deserve to be reexamined. Metal-on-metal

bearings were originally used in total hip replacements in the 1950s, with limited

success. Complete ceramic bearings were seen as early as the 1970s, but fracture

of the ceramic components and high wear were attributed to problems in proc-

essing that resulted in defects and inadequate grain sizes (Plenk, et al., 1992;

Sedel, et al., 1991). In time, they were almost completely replaced in the clinical

world by metal–polymer combinations. However, their revival in the past few

1180 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Fig. 6 Schematic diagram of an artery showing the three tissue layers:

adventitia (outer), media (middle), and intima (inner).

years, spurred on by developments in materials engineering, indicates that the

original developers had the right idea—they were just a few years ahead of their

time.

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES

When blood vessels are damaged through injury or disease, they often must be

replaced or bypassed in order to maintain adequate blood ﬂow to and from the

regions of the body. Disease-induced damage, such as atherosclerosis and an-

eurisms, occurs more often in arteries than in veins, due in large part to the

higher working pressure of the blood within these vessels. Injury can occur to

any blood vessel; however, collateral circulation typically eliminates the need to

replace small veins, and the low venous return pressure provides an environment

in the larger veins that is much more conducive to traditional surgical repair or

autograft use. As a result, this section will focus on the selection of materials

for the development of arterial prostheses.

Arteries are three-layer hollow tubes (Fig. 6) composed of a combination of

elastin, collagen, and smooth muscle—with the proportions of each component

dependent on the size of the artery. The size of an artery also varies substantially

along the arterial tree, from a typical diameter of about 4 cm for the aorta in an

adult to a diameter of less than 1 mm for arterioles. The age and size of the

individual will also affect the dimensions of the arteries. Normal arterial pressure

is approximately 120 mmHg during the systolic phase (contraction) of the car-

diac cycle, decreasing to approximately 70 mmHg during the diastolic phase

(relaxation). However, pressures can rise as high as 200 mmHg or more in

individuals with either transient or chronic hypertension, and hypertension itself

is a risk factor for a number of the pathological processes that may require

arterial replacement or bypass.

Elastic arteries, such as the aorta, are expected to expand during systole as

blood is forced into them by the heart. The elastic recoil of these vessels, due

in large part to the high amount of elastin in the wall, acts to maintain arterial

pressure during the relaxation phase of the cardiac cycle, further transporting

blood away from the heart. Distributing arteries, which (as their name indicates)

distribute blood from the larger, elastic arteries to the various regions of the

body, have more collagen than elastin in their walls. However, the physiologi-

cally important component of these vessel walls is the smooth muscle, which

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES 1181

allows the distributing arteries to adjust their diameter and therefore their resis-

tance. By controlling the arterial resistance to blood ﬂow in various regions of

the body through vasodilation and constriction, blood can be directed to regions

that have the highest metabolic demand—the muscles during exercise or the

gastrointestinal system during digestion.

Vascular grafts were ﬁrst introduced in the early twentieth century as solid

tubular structures. Fabric grafts were developed in the 1950s and provide the

basis for the grafts that are currently in use or under development. An arterial

graft or arterial prosthesis has one obvious function—to transport blood from

vessels that are proximal to the graft to those that are distal to the graft. In

addition, the vasodilation/constriction and elastic recoil capabilities of the var-

ious types of arteries assist the heart with transport and direction of blood to the

diverse tissues of the body. However, there are additional functional constraints

that are equally as important. First, the prosthesis must maintain hemostasis—

not allow blood to leak from the artery into the surrounding tissues. Second,

interaction of the blood cells with the vessel must not act to initiate coagulation,

which could result in thrombus formation and embolisms, nor hemolysis, the

destruction of red blood cells. These functional constraints will be considered

in the discussion on biocompatibility.

3.1 Function

Blood Transport

To transport blood between two connecting vascular segments, a tubular structure

is required. However, given modern material processing techniques, this require-

ment does not itself limit the selection of materials substantially. The graft must

also, however, be connected to the ends of the remaining vascular segments in

some way, either through ligatures or sutures.

Using a maximum arterial pressure of 200 mmHg (30 kPa), the circumfer-

ential and axial wall stresses within a graft (diameter of 1 cm, thickness of 0.5

mm) can be approximated to reach 270 and 135 kPa, respectively. This is well

within the mechanical limits of most undegraded artiﬁcial materials; however, it

may become a substantial constraint for vessels engineered from natural mate-

rials in the future. In addition, the physiological environment, combined with

the cyclic loading seen by the vessel due to the normal vascular pressure vari-

ations, will affect the properties of the graft material. As discussed above for

hip replacement, corrosion, absorption, and leaching processes that occur when

a material is placed in vivo will all negatively affect the ultimate strength of a

material.

The expansion and recoil observed in elastic arteries must be taken into con-

sideration when selecting a material for grafts at these locations. It is reasonable

to assume that a relatively short length of the graft does not itself need to

dynamically change dimensions to maintain blood ﬂow through the segment of

vasculature. However, any mismatch in the behavior of the graft to the connected

vessels will result in substantial stresses on the ligatures, sutures, or the vessel

itself. In certain applications, the graft may replace a substantial portion of an

artery, such as in the descending aorta. In this case, elastic recoil in the graft

will play an important role in maintaining the velocity of blood ﬂow through

1182 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

Fig. 7 Schematic diagram showing the characteristic nonlinear force–deformation relationship

of an artery. Initial region of high compliance is characteristic of the elastin component in the

walls, while the later, low-compliance region exempliﬁes the deformation of the collagen ﬁbers.

the distal vessels, and the pulsatile ﬂow may be important to maintaining func-

tion in organs supplied by the arterial tree (Mergerman and Abbott, 1983). The

same philosophy would apply to distributing arteries. At the current time, no

artiﬁcial material exists that would respond to the physiologic control mecha-

nisms in order to dilate and constrict in conjunction with surrounding vessels.

However, a graft that is less compliant than the attached vessel would result in

substantial stress concentrations at the junction between the artiﬁcial and natural

materials. This elevation in stress could lead to failure of the sutures or ligatures

connecting the structure, as well as accelerate fatigue processes within the graft

ends (Wilkerson and Zalina, 1994). In addition, a less compliant graft can also

result in a stenosis in the vascular tree during systole, when the proximal seg-

ment of artery expands under pressure, but the blood then encounters the length

of reduced diameter graft that acts to retard ﬂow (Herring, 1983). These issues,

along with the fact that blood vessels experience signiﬁcant deformation due to

general body motion, essentially dictate that a soft, ﬂexible material be used for

the graft.

An ideal graft would exactly match the compliance of the attached blood

vessel, allowing it to expand and recoil to the same extent as the natural struc-

ture. This would require a nonlinear response, with an initial region of highly

compliant stretch, attributed to the unkinking and realigning of the collagen

ﬁbres and a simultaneous stretching of the elastin. After the collagen in the

vessel wall is straightened, its higher elastic modulus dominates the behavior of

the vessel, and the compliance is signiﬁcantly reduced. It is this behavior (Fig.

7) that is thought to limit the overexpansion of arteries during acute episodes of

elevated blood pressure. The compliance of each segment of artery therefore

depends on the proportion of collagen, elastin, and smooth muscle (which has

a negligible effect on compliance in its passive state) within the vessel wall.

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES 1183

Typical values for compliance (in units of percent radial change per millimeter

of mercury) for cadaveric human arteries range between 10.9 at 60 mmHg and

3.8 at 150 mmHg (Mergerman and Abbott, 1983).

3.2 Biocompatibility

The primary functional need of a blood vessel—transport of blood—can easily

be met through general implant design. However, due to the delicate nature of

blood cells and the ease at which the clotting cascade can be initiated, biocom-

patibility issues place substantial limitations on material selection for this ap-

plication. The natural vessel provides an optimal environment for blood ﬂow,

and the mimicking or replacing of its intimal layer is one of the underlying ideas

in work to improve biocompatibility in vascular grafts.

Neointima Formation

Natural blood vessels consist of three layers: the outer adventicia, the middle

media, and the inner intima. The intima has a layer of endothelial cells that

contact the moving blood in order to minimize damage to the cellular elements

of the ﬂuid. Some artiﬁcial materials that are implanted into the body as vascular

grafts have been shown to develop a neointima—ﬁbrin and ﬁbroblasts that de-

velop as a lining on the inside surface of the tube (Herring, 1983).

The rate of neointima formation and its constituents may depend on the ma-

terial used in the vascular graft. A carbon–ceramic composite with a large degree

of surface porosity showed almost instantaneous (

⬍10 s) development of a ﬁbrin

mesh with limited platelet adhesion (Chignier et al., 1987). This developed into

a 5- to 7-cell thick layer of ﬁbroblasts, collagen, and elastin by 15 days, with

endothelial-like cells present and fully developed by 2 months. Herring (1983)

provides a good description of the ‘‘healing’’ of vascular grafts and the devel-

opment of the neointima. In most human vessels, complete endothelialization

occurs only in the vicinity of the anastomoses with the natural vessels, as cells

grow in from the ends of the connected tissue. To expand the area of vessel that

experiences complete healing—such that the lining resembles that of a natural,

healthy vessel—research is being conducted into seeding the vessel wall with

endothelial cells (Consigny, 2000; Herring, 1983). This can either be done during

the preclotting process, in the operating room, using a small number of cells

taken from a vein that is exposed during the surgical procedure, or preoperatively

using cells from the jugular vein that are cultured to provide a greater volume

for the seeding process. In both cases, the use of autologous (patient speciﬁc)

cells is important to eliminate issues of rejection.

The geometry of the vessel also affects neointima formation. The initial in-

teraction between the blood and the graft is the formation of a clot, or thrombus,

on the inner surface of the vessel. The thickness of the implant wall is directly

proportional to the thickness of the thrombus formed on its surface (Park and

Lakes, 1992). As the thrombus must be remodeled to form the mature neointima,

a smaller clot results in faster organization of the neointima. Neointima forma-

tion must be limited, however, to prevent vessel occlusion. This is one of the

issues with small diameter vascular grafts, where the ﬁbrin layer may continue

to grow past 1 mm in thickness (Wilkerson and Zalina, 1994). Finally, it is

important to ensure that the neointimal layer that develops maintains its integrity

1184 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

and does not change in such a way, either through mineralization or adsorption

of additional organic components, that it no longer promotes the continued pa-

tency of the vessel (Hufnagel, 1983).

Hemostasis

For vessels to adequately transport blood, the blood must remain within the

lumen of the vessel and not leak into the surrounding tissues in any great

amount. The obvious answer to this problem is to employ solid materials for

vascular grafts, nonporous structures that will adequately contain the blood.

However, the development of a neointima within a vessel provides a second

design option, as it will act to seal a porous blood vessel against blood leakage.

However, the time course for neointima formation is not short enough to provide

the immediate seal needed during surgery and postoperatively. Therefore, many

vascular grafts are preclotted in the operating room before they are implanted

by exposing both the inner and outer surfaces to the patient’s blood. This pro-

vides an initial surface that serves as the basis for neointima formation.

Hemolysis

Hemolysis, or the damage and destruction of red blood cells, occurs continually

within the body. The cells are constantly replaced by new erythrocytes produced

by the bone marrow. However, chronic damage to these cells and the release of

their cellular contents into the plasma of the bloodstream can result in anemia,

kidney failure, and other toxic reactions (Hershko, et al., 1998). Hemolysis due

to nonphysiologic mechanisms typically occurs due to high shear stresses. A

shear stress as low as 400 Pa can damage or rupture a red blood cell that comes

in contact with another solid surface (Sallam and Hwang, 1984). Turbulent ﬂow

can also result in sufﬁcient shear stresses to cause hemolysis. Therefore, two

goals of vascular graft design are to minimize the contact between red blood

cells and the graft surfaces and minimize the turbulence that may occur at

branches or divisions of blood vessels. The latter falls within the domain of ﬂuid

mechanics, rather than biomaterial selection.

To minimize the contact between blood cells and graft walls that may lead

to hemolysis, it is desirable to form a natural tissue layer between the two. Graft

materials with high surface tensions tend to initially attract platelets and ﬁbrin

molecules, which aggregate and form the natural boundary that is then remod-

eled to become the neointima. The elements of the tissue layer also exhibit a

slight negative electrical charge (Collins, 1983), which may act to repel the

negatively charged red blood cells, therefore additionally reducing the contact

between the cells and the vessel wall.

Coagulation

Blood clotting, or coagulation, is a necessary physiologic process that allows

the body to heal and maintain its blood pressure through hemostasis. However,

the formation of stationary clots within blood vessels (thromboses) and the

movement of those clots with the blood ﬂow (emboli) can result in vascular

blockage, tissue damage, and even death. Therefore, it is important to select

materials for graft use that will minimize the initiation of the clotting cascade.

The effect of a material on blood clotting is not easy to assess, particularly

as it is difﬁcult to separate the effect of the material from the natural physiologic

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES 1185

process of coagulation. Blood will also behave slightly differently in vivo than

it will in vitro and, like other tissues, can differ between individuals. Therefore,

an in vitro coagulation test with blood from a dog may not be fully indicative

of how that material will behave when implanted into a human. Standards have

been developed, however, for coagulation tests (Bruck, 1980), and these progress

from static, in vitro tests to dynamic, in vivo tests.

Surface tension, surface charge, and surface roughness are properties of a

material that will affect the rate and amount of coagulation that takes place when

it is in contact with blood. Polymeric materials tend to adsorb a mixture of

proteins to their surfaces in the ﬁrst 30–60 s of contact with physiologic ﬂuid

(Baier, 1975), the composition of which depends on the polar or nonpolar nature

of the polymer (Herring, 1983). Adsorbtion and activation of key molecules from

blood, including Hageman factor, factor XI, and others, will trigger the clotting

cascade (Forbes, 1993). Rougher surfaces, including crimped grafts, have been

shown to increase the rate of coagulation that occurs when in contact with blood

(Collins, 1983). This is probably due to the larger surface area that can come in

contact with the blood. A rough surface may be desirable, however, in order to

promote preclotting on the surface of a porous graft. In the case of crimping,

the process also prevents kinking of the vessels during surgery or prolonged

implantation, which itself can lead to occlusion. Surface charge can help to

minimize the contact between the graft and the blood elements. The formed

elements of the blood—red and white blood cells and platelets—have been

shown to have a negative surface charge. As a result, a vascular surface with a

slight negative charge, through the presence of either a neointima or a naturally

or artiﬁcially induced surface charge, will act to repel the blood cells and plate-

lets away from the vascular wall (Collins, 1983). When platelets do not come

in contact with the vascular wall or foreign bodies, they are less prone to initiate

the clotting cascade.

The natural lining of blood vessels possesses unique properties that cannot

be easily mimicked. It was originally believed that the smooth surface and the

negative surface charge were the properties of an endothelial lining that pre-

vented clot formation. However, studies have shown this to be more complex.

In particular, the presence of endothelial cells allows for the secretion of para-

crine agents that act to break down small thromboses and interfere with clot

formation (Bruck, 1980). Smooth muscle and ﬁbroblasts did not exhibit the same

function, and in some cases precipitated platelet activation. Thus, as was the

case in bone’s resistance to fatigue, the living function of vascular tissue cannot

be fully replaced using current technologies and artiﬁcial materials.

To reduce clotting in artiﬁcial materials, several approaches have been taken.

Heparin, a negatively charged polysaccharide that is commonly used to prevent

clottting in many clinical applications, has been coated on implants. Using the

same logic, anionic radicals have been included in an artiﬁcial material to pro-

duce the negative surface charge that only naturally can occur in polymers.

Taking a different approach, materials with low surface tensions have been pro-

posed, as they are less likely to attract the formed blood elements to the material

surface and initiate a clotting cascade. This last tact is, of course, in contradiction

to the suggested use of high surface tension materials to minimize hemolysis!

This latest dilemma is indicative of many decisions that must be made in

selecting a material for use in biomedical applications. No material—besides

1186 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

possibly the original tissue—will perfectly match all of the goals and constraints

of the design. It is generally necessary to weigh the beneﬁts and drawbacks of

each material that makes the short list and to select the one that shows the best

balance. While this is frustrating from a design aspect, it is what spurs material

development and allows for the continuing work in the area of material synthesis

and implant design.

Property Degradation

Once a material is selected that does not seem to have a signiﬁcant, adverse

effect on the blood itself, it is necessary to evaluate how the properties of the

graft are affected during the expected contact with the physiological environ-

ment. As discussed above for orthopedic implants, polymers can react to the

physiologic environment by either leaching smaller molecules into the surround-

ing tissue or absorbing water or other materials from the tissue. In the case of

vascular grafts, the constant ﬂow of blood past the implant surface will eliminate

the possibility to develop an equilibrium between the molecule concentrations

inside and outside of the implant material. As an example, nylon has been found

to absorb water from the surrounding environment when implanted (Edwards,

1983). The water molecules act as plasticizers, which reduce the cohesion and

bonding between the chains of the polymer. The ﬁnal result is a reduction in

strength.

Whether property degradation will be a determining factor in material selec-

tion for vascular implants depends largely on the planned time of implantation

for the device. For a graft that will permanently replace a section of blood vessel,

it is extremely important to maintain the tissue properties at an appropriate level.

However, many blood contacting applications involve short-term use—for in-

stance intravenous catheters for administering blood or pharmaceuticals. In these

cases, the length of intended use is typically less than a few days, and the

catheter can be relatively easily replaced if necessary. Therefore, for short-term

implants, the prevention of blood clot formation and blood cell damage, along

with ease of use and cost, become the determining factors in material selection.

3.3 Current Material Selection

Since the 1950s, polymer fabrics have been the primary material used for long-

term vascular grafts. Nylon was introduced in 1955, but was withdrawn from

use after the occurrence of aneurisms within the grafts indicated a loss of me-

chanical integrity in vivo (Edwards, 1983). Teﬂon (PTFE, Gore-Tex) and poly-

ethylene terephthalate (PET, Dacron) have been shown to have acceptable

amounts of property degradation when implanted for 20 years or more (Snyder,

1983). Table 5 shows results on the loss of strength in various polymers that

were implanted for 100 days. The Dacron showed an initial drop in strength that

then stablized at an acceptable level.

Typical vascular grafts are constructed of a woven or knitted fabric and are

crimped, to both prevent kinking and to allow for longitudinal expansion (Fig.

8). The response of the graft in vivo will depend not only on the constituent

material selected, but also on the weave of the fabric and on its processing

(Sawye, et al., 1983). Grafts are typically tested to validate their tensile or burst-

ing strength, and the values obtained for the constructed vessel will differ from

those of the bulk material due to the knitted or woven nature of the fabric.

3 BLOOD-CONTACTING BIOMATERIALS: VASCULAR PROSTHESES 1187

Table 5 Average Changes in Tensile

Properties of Synthetic Grafts Implanted

for 100 Days

Material Loss of Strength (%)

Nylon ⫺81

Orlon

⫺6.9

Dacron

⫺10.1

Teﬂon

⫹3.2

Source: Data from (Edwards, 1983)

Fig. 8 Left and right: Examples of fabric arterial grafts, illustrating the crimped structure that

prevents kinking in vivo. Center: Extruded PTFE graft lined with pyrolytic carbon.

Pyrolitic or LTI (low-temperature isotropic) carbon has been found to have

excellent anticoagulation properties (Bruck, et al., 1973), without the need for

heparin coating of the material. The original LTI formulation required a solid

substrate and so was more conducive for use on shunts or leaﬂets of heart valves.

A newer, ultra-low-temperature iostropic carbon (ULTI) can be vapor deposited

onto fabrics as well (Bruck, 1980), allowing for its use in regions that require

ﬂexibility (Fig. 8). ULTI carbon, which can be deposited with a thickness of

less than 1

␮

m, has been shown not to affect the compliance of the underlying

fabric (Sharp, 1983), indicating that it will maintain its ﬂexibility when deposited

on a fabric graft.

For short-term implants, generally in the form of catheters, silicone rubber

has become the standard. Its optical translucence makes it easy to monitor ﬂuid

transfer, it is easy to use, and it is highly biocompatible.

Fabric grafts and silicone tubes, however, are not adequate for the small-vessel

replacement that may be required during reconstructive surgery. As a result,

1188 SELECTION OF MATERIALS FOR BIOMEDICAL APPLICATIONS

tissue engineering techniques are currently being employed to construct small-

diameter vessels from natural materials, such as chitosan (Chupa, et al., 2000).

Research such as this is working to expand the possibilities for material selec-

tion, not only for vascular prostheses but also for all biomedical implants.

4 SPACE-FILLING BIOMATERIALS: BREAST IMPLANTS

Implants designed to ﬁll voids where tissue has either been removed or destroyed

are commonly used in reconstructive surgery. Hard tissue, soft tissue, and com-

bination materials have been utilized to replace or augment cranial defects, the

loss of an ear, nose, or eye due to trauma or disease, and congenital facial

abnormalities, among others. Perhaps the most familiar and infamous space-

ﬁlling implant was, and is, the breast implant. Originally designed for recon-

structive surgery following mastectomies, breast implants became increasingly

popular for cosmetic enhancement during the 1970s and 1980s. The question of

whether the implants were the cause of the illnesses and disabilities reported in

the 1990s has never been answered to everyone’s satisfaction, and probably

never will be. It is an example of a biocompatibility issue that, if true, went

unnoticed for many years. It also serves notice to all biomedical implant de-

signers that because an implant does not indicate any problems after 5 years

does not mean that monitoring of potential complications should stop.

Breast implants were designed with one function in mind—to replace tissue

that had either been removed through surgery or that an individual felt nature

had left lacking. There was no other physiologic role for the implant, which

made its design much simpler. The ﬁrst breast augmentation utilized injections

of silicone, parafﬁn wax, or bees wax directly into the tissue surrounding the

breast. This method was banned by the Food and Drug Administration (FDA)

in the 1960s as the injected material was seen to migrate and lose its shape. In

addition, because of the large contacting surface area between the injected ma-

terial and surrounding tissue, adverse tissue reactions were seen (Frisch, 1983).

Since that time, breast implants have utilized conﬁned volumes of materials

(Figure 9).

4.1 Function

Space Filler

As stated above, the single role of a breast implant is to ﬁll up a given volume

of space within the body. However, cosmetic appearance and ‘‘feel’’ have also

dictated much of the development of the implants. The density of the implant

should be similar to or less than the surrounding tissue, so that tissue damage

does not occur due to increased weight. The consistency of the material should

also be somewhat similar to the composite of fatty and connective tissue that it

has replaced. Thus, metals, ceramics, and solid pieces of polymer would be

inappropriate.

4.2 Biocompatibility

Capsule Formation

No artiﬁcial material that is implanted into the body will be ignored completely.

The degree of the reaction from the immune system will depend on a number