Bio Materials

TUTORIAL

BIOMATERIALS

Oleh:

dr. Hidayat

Pembimbing:

dr. Tjuk Risantoso,SpB,SpOT

Orthopaedi & Traumatologi Fakultas Kedokteran Universitas Brawijaya Malang

2015

Definition

A biomaterial is a nonviable material used in a medical device, intended to interact with

biological systems. Defined by their application NOT chemical make-up.

The term biomaterials refers to synthetic and treated natural materials

that are used to replace or augment tissue and organ function.

Biomaterials must meet several criteria to perform successfully. They

must be biocompatible, or able to function in vivo without eliciting an

intolerable response in the body either locally or systemically; resistant to

corrosion and degradation, meaning that the body environment must not

adversely affect material performance; and possess adequate mechanical

properties, an especially important criterion for those biomaterials used in

devices intended to replace or reinforce load-bearing skeletal structures.

In addition, orthopaedic biomaterials intended for total joint replacement

must possess adequate wear resistance to maintain proper joint function

and to minimize biocompatibility problems caused by biologic reactions to

particulate debris. They must be capable of reproducible fabrication to the

highest standards of quality control and, of course, at a reasonable cost.

Biomaterials that meet these criteria are fundamental to the practice of

orthopaedic surgery. They have been used successfully to develop

devices for internal fixation of fractures, osteotomies and arthrodeses,

wound closure, softtissue reconstruction, and total joint arthroplasty that

have advanced significantly the treatment of musculoskeletal diseases.

But at the same time, limitations in biomaterial performance, failure of the

implant designer to understand the limitations, or inappropriate

application of the technology often are directly related to clinical failure.

Therefore, an understanding of the physical and chemical properties of

orthopaedic biomaterials is an important consideration in selecting and

using implant devices and in providing a realistic expectation of clinical

performance. This chapter will discuss how a material’s physical

properties result from its chemical composition and structure, and

fundamentals of mechanical performance in terms of the important

material properties that determine the behavior of implant structures.

Common metallic, polymeric, and ceramic orthopaedic biomaterials are

described in terms of their molecular structures, microstructures,

composition, and properties, and how their properties are influenced by

processing and manufacturing variables is discussed as well.

Characteristics of Biomaterials

Physical Requirements

Hard Materials.

Flexible Material.

Chemical Requirements

Must not react with any tissue in the body.

Must be non-toxic to the body.

Long-term replacement must not be biodegradable.

History

More than 2000 years ago, Romans, Chinese, and Aztec’s used gold in dentistry.

Turn of century, synthetic implants become available.

1937 Poly(methyl methacrylate) (PMMA) introduced in dentistry.

1958, Rob suggests Dacron Fabrics can be used to fabricate an arterial prosthetic.

1960 Charnley uses PMMA, ultrahigh-molecular-weight polyethylend, and stainless

steal for total hip replacement.

Late 1960 – early 1970’s biomaterial field solidified.

1975 Society for Biomaterials formed.

First Generation Implants

“ad hoc” implants

specified by physicians using common and borrowed materials

most successes were accidental rather than by design

Examples — Second generation implants

titanium alloy dental and orthopaedic implants

cobalt-chromium-molybdinum orthopaedic implants

UHMW polyethylene bearing surfaces for total joint replacements

heart valves and pacemakers

Artificial Hip Joints

http://www.totaljoints.info/Hip.jpg

MED

BOSC II Surakarta 2014

Third generation implants

bioengineered implants using bioengineered materials

few examples on the market

some modified and new polymeric devices

many under development

Example - Third generation implants

tissue engineered implants designed to regrow rather than replace tissues

Integra LifeSciences artificial skin

Genzyme cartilage cell procedure

some resorbable bone repair cements

genetically engineered “biological” components (Genetics Institute and Creative

Biomolecules BMPs)

Grow cells in culture.

Apparatus for handling proteins in the laboratory.

Devices to regulate fertility in cattle.

Aquaculture of oysters.

Cell-silicon “Biochip”.

Metals

Semiconductor Materials

Ceramics

Polymers

Synthetic BIOMATERIALS

Orthopedic screws/fixation

Dental Implants

Dental Implants

Heart valves

Bone replacements

BiosensorsImplanta

ble Microelectrodes

Skin/cartilage

Drug Delivery Devices

Ocular implants M

ED BO

SC II Surakarta 2014

Examples of Biomaterial Applications

Metals

Metallic alloys have found widespread use in orthopaedic surgery. Alloys are metals

composed of mixtures or solutions of metallic and nonmetallic elements. The combination of

elements is used to impart the high strength, ductility, and elastic modulus, the corrosion

resistance, and the biocompatibility required for load-bearing applications, such as fracture

fixation devices and implant components for total joint arthroplasty. The 3 common alloys

used in orthopaedics are stainless steel, cobalt chromium alloy, and titanium alloy. None of

these alloys were developed specifically for orthopaedic or biomedical applications.

Instead,their proven strength and corrosion resistance in the aerospace, marine, and chemical

industries have led them to be adopted for implant use. Their specific properties can be

understood from their molecular structure, microstructures, and composition.

Stainless Steel

The most common form of stainless steel used in orthopaedic applications is 316L, grade 2,

designated by the American Society for Testing and Materials (ASTM) specification F138.

The numeric designation “316” places the alloy within the so-called austenitic stainless

steels; the “L” denotes low carbon concentration (typically below 0.03 wt%). As with all

steels, 316L stainless steel is an alloy of iron and carbon. The other major alloying elements

include chromium, nickel, and molybdenum, with minor amounts of manganese,

phosphorous, sulfur, and silicon. The alloying elements affect the microstructure and hence

the mechanical and corrosion properties of the steel. Chromium in the microstructure forms a

strongly adherent oxide (Cr2O3) on the surfaces of the metal that are exposed to the

environment, thus providing corrosion resistance by forming a passive layer between the

environment and the bulk material. Stainless steel devices are passivated by immersion in a

strong nitric acid bath as part of the manufacturing process to assure the creation of the oxide

layer. The creation of a “passive” oxide layer limits the rate of electrochemical corrosion by

about a thousand to a million times compared to the rate of corrosion in the absence of the

oxide. Most of the metallic alloys are highly reactive with oxygen, so that the oxide layer

forms naturally as the base metal is exposed to the atmosphere. Standardized methods (such

as the nitric acid bath) are used to enhance the layer and ensure appropriate resistance to

corrosion. Though chromium provides the “stainless” quality to the steel, it also stabilizes the

ferritic, body-centered cubic phase that is weaker than the face-centered cubic austenitic

phase. Molybdenum, added to provide additional corrosion resistance, and silicon, added

with manganese to aid in the manufacturing process, also stabilize the ferritic phase. To offset

this tendency, nickel is added to stabilize the austenitic phase and thus assure an

appropriately strong microstructure.

Carbon concentration must be kept low in 316L stainless steel to maintain corrosion

resistance. At higher carbon concentrations, there is a tendency for the carbon to combine

with the chromium to form a brittle carbide that robs the microstructure of much of the

chromium and that tends to segregate to the grain boundaries, significantly weakening the

material by making it prone to corrosionrelated fracture. Such a condition, called

sensitization, has been directly responsible for mechanical failures of orthopaedic implants

made from stainless steels in which carbon content has been too high. ASTM specifications

for 316L stainless steel call for an austenitic microstructure free of carbides or inclusions that

might remain from the steel-making process (and that can reduce corrosion resistance). The

recommended grain size is small (about 100 microns in any dimension) to assure adequate

strength for orthopaedic applications. Grain size can be controlled by the solidification

process and by postsolidification heat treatments and cold working of the material.

Stainless steel is typically cold-worked by about 30% for orthopaedic applications.

Mechanical properties of 316L stainless steel are provided in Table 2 in both the annealed

(not cold-worked) and 30% cold-worked condition. A potential disadvantage of stainless steel

in implant applications is its susceptibility to crevice and stress corrosion. In any corrosion

process, there are 2 reactions, an anodic reaction in which the metal is oxidized to its ionic

form (M → Mn+ + n electrons) and a cathodic reaction in which the electrons are consumed

(in an aqueous solution with dissolved oxygen, O2 + 2H2O + 4e →4OH–). These reactions

could initially be progressing at an even rate over the surface of a stainless steel implant, such

as a bone plate or a bone screw. But as the reactions progress in the crevice between the

underside of the head of the screw and the countersunk area of the plate, the crevice becomes

depleted of oxygen. The anodic reaction continues in the crevice while the remainder of the

plate and screw undergo the cathodic reaction. The oxygen concentration is not readily

replenished by the fluids outside of the crevice, though smaller chlorine ions flow into the

crevice, drawn there by the metal ions being released by the anodic reaction.

The crevice region decreases in pH, causing accelerated metal oxidation. Stress corrosion

cracking results when the combination of an applied stress and a corrosive environment lead

to mechanical failure of the material, even though the environment or load in and of itself

would be insufficient to cause failure. Stress corrosion cracking has been shown to occur

even under low levels of constant stress, such as might occur in an implant with residual

stresses. Crack initiation is accelerated by the corrosion process, as is the subsequent crack

growth that occurs under the applied stress. Because of concerns about corrosion and

subsequent long-term biocompatibility, stainless steel has been used primarily in fracture and

spinal fixation applications. These applications often allow removal of the device or require

strength only until healing occurs. Permanent implants, such as femoral components of the

Charnley design of hip replacements, have also been made from stainless steel, demonstrating

that stainless steel can be used safely even in these high-demand applications.

Cobalt-Chromium Alloys

Cobalt-chromium alloys include compositions intended to be manufactured by casting

(ASTM F75 alloy) and by forging (ASTM F799 alloy), as well as alloy compositions that

obtain excellent mechanical properties through cold working (ASTM F90 and F562). All of

these alloys are primarily cobalt with significant amounts of chromium added for corrosion

resistance. As with stainless steel, the chromium forms a strongly adherent oxide film that

provides a passive layer shielding the bulk material from the environment. The F75 and F90

alloys contain about 60% cobalt with about 28% chromium. The F799 and F562 alloys have

less cobalt and chromium, and in their place have large amounts of other alloying elements

(about 15% tungsten in F799 and about 35% Ni in the F562).

The alloys display a range of mechanical properties that can be understood from the

processes and the resulting microstructures used to fabricate devices from the materials. The

F75 alloy, for example, has commonly been used for investment (or so-called lost wax)

casting. Waxmolds of near-final dimensions of devices such as total hip femoral stems are

coated with a ceramic slurry. The slurry is fired in a furnace (and the wax is lost as it melts

away from the inside of the ceramic mold). Molten F75 alloy is poured or pressurized into the

molds and allowed to solidify. The ceramic mold is broken away from the underlying metal

part, which can then be finished into the final device. Quality control can be a problem during

the casting process. If solidification proceeds too slowly, grains have ample time to grow

quite large, thus significantly diminishing the material’s strength (Fig. 12). If solidification

proceeds too quickly, air from inside the mold and gases that are released during the

solidification process can become entrapped in the microstructure, causing undesirable stress

concentrations that can cause premature failure. Finally, if cooling conditions are not ideal,

carbides that naturally occur within the alloy’s microstructure segregate to too great a degree,

which can weaken the material, reduce ductility, and decrease corrosion resistance. To

overcome these problems, the alloy can be fabricated by powder metallurgy. Fine powder of

the alloy is compacted and sintered together to form a near net shape. The shape is then

forged under pressure and heated into the final shape. The resulting microstructure has a

smaller grain size and more evenly distributed carbides than the cast alloy, leading to

improved properties F75 alloy is used to fabricate porous coatings for biologic fixation of

orthopaedic implants. The resulting properties of the porous-coated device will depend on the

microstructure of the substrate metal and the porous beads, as well as the thermal sintering

process that is used to connect the two. Sintering involves very high temperatures (near the

1225°C melting temperature), which can significantly decrease the fatigue strength of the

substrate material.

Together with the stress concentrations caused at the attachment points with the porous

coating, the result is a fatigue strength of only about 200 MPa, even after additional thermal

treatments are used to restore some of the strength. This strength is well below that achieved

for other cobalt alloys that are not porous-coated. The forging alloy, F799, possesses

mechanical properties that are superior to that of the cast alloy. Hot forging effectively

reduces grain size, “heals” pores through the combination of pressure and heat, and breaks up

the carbides into an even distribution. The thermomechanical forging operation also induces

an additional microstructural phase that contributes to the improved properties.

The F90 and F562 alloys obtain substantial mechanical properties through more than 40%

cold-working. The tungsten addition in F90 improves machinability and fabrication via cold-

working. The cold working of F562 alloy provides additional energy for the transformation of

some of the face-centered cubic phase into a hexagonal phase that emerges as fine platelets

throughout the microstructure. The combination of a very fine grain size (the facecentered

cubic grains are less than 0.1μ in any dimension) and the dispersed platelets impede plastic

deformation, strengthening the material. In addition, the material can be thermally treated to

precipitate a uniform distribution of very fine cobalt-molybdenum (Co3Mo) precipitates that

act to further strengthen the material. The result is among the strongest of the orthopaedic

implant alloys.

The ease of fabrication and the range of properties available for cobalt alloys make them deal

for a wide range of orthopaedic applications, including all metallic components of all joint

replacements as well as fracture fixation devices. The chromium content of these alloys

provide excellent corrosion resistance, with superior resistance to crevice corrosion than

stainless steel. Long-term clinical use has proved that these alloys also have exceptional

biocompatibility in bulk form

Titanium and Titanium Alloys

Titanium and its alloys are of particular interest for biomedical applications because of their

outstanding biocompatibility and corrosion resistance. Their corrosion resistance, provided

by an adherent passive layer of titanium oxide (TiO2), significantly exceeds that of stainless

steel and the cobalt alloys. Uniform corrosion even in saline solutions is extremely limited,

and resistance to pitting and intergranular and crevice corrosion is excellent. Experimental

studies in animal models and long-term clinical use in humans confirm truly superior

biocompatibility. Furthermore, the oxide surfaces of titanium and its alloys are well tolerated

in contact with bone, becoming osseointegrated with little evidence of a fibrous layer

between bone and implant.

CP-titanium (ASTM F67) is used more extensively in dental implants, but is used in

orthopaedic surgery primarily in the form of wire mesh for porous coatings that is sintered

onto titanium alloy joint replacement components. The properties of CP-titanium depend on

the amount of oxygen contained in the metal. At small concentrations, increased oxygen

content improves the mechanical properties. Grade IV CP-titanium, for example, with an

oxygen concentration of 0.40 wt% has a yield strength of about 485 MPa, while grade I with

an oxygen concentration of 0.18 wt% has a yield strength of only 170 MPa. The CP-titanium

microstructure consists of grains of a single, hexagonal close-packed phase, and the material

can be cold-worked. Additional strengthening of the microstructure comes from interstitial

solid solution strengthening, in which atoms of oxygen, carbon, and particularly nitrogen

harden the material by being encased in the interstices of the crystalline, hexagonal

arrangement of titanium atoms.

The most common form of titanium used in orthopaedic applications is titanium-aluminum-

vanadium alloy (ASTM F-136). The primary alloying elements, aluminum and vanadium, are

limited to 5.5 wt% to 6.5 wt% and 3.5 wt% to 4.5 wt%, respectively (Fig. 13), so that the

alloy is often called Ti-6Al-4V or simply Ti-6-4. Developed by the aerospace industry as a

high strength-to-weight ratio material, the alloy is used in orthopaedic implants in the extra

low interstitials form, in which the oxygen concentration is kept very low to avoid

embrittlement and to maximize strength and ductility. The microstructure of Ti-6Al-4V

contains 2- phase grains, the alpha phase being a hexagonal-close packed phase that is

stabilized by the aluminum alloying element and the body-centered cubic beta phase

stabilized by the vanadium. The distribution and amount of the phases dictate the material’s

properties and can be altered by prior thermal treatments. The alloy can also be mechanically

worked to alter its properties. Typically, the microstructure is a fine-grained alpha phase with

the beta phase present as isolated particles that precipitate at the grain boundaries; this

microstructure possesses excellent fatigue resistance compared to other forms of titanium

alloy microstructures.

The mechanical properties of Ti-6Al-4V are more than adequate for most orthopaedic

applications The elastic modulus for the alloy is about half that of stainless steel and the

cobalt alloys, making the alloy an ideal candidate for lowering the structural stiffness of a

device without changing its shape. For example, the axial, bending, and torsional stiffnesses

of a bone plate fabricated from titanium alloy will be half that of a bone plate of the same size

and shape made from stainless steel or cobalt alloy. Thus, the severity of stress shielding

when the plate is rigidly attached to the bone (so that the bone and the plate share

load) would be less for the titanium alloy plate. This mechanical consideration has led to the

use of titanium alloy in fracture and spinal fixation devices, including plates, nails, and

screws. The same consideration has led to the use of titanium alloy in stems for total joint

replacements.

A disadvantageous trait of titanium alloy is its notch sensitivity. A stress concentration, such

as a notch or scratch, on the surface of a titanium alloy implant significantly reduces the

fatigue life of the part. The same is true for the type of stress concentrations that occur when

a porous coating is applied to the surface of a titanium alloy total joint component. The

severe changes in geometry that result from the sharp angles that are created wherever the

coating is sintered to the substrate act as points of stress concentration. Therefore, care must

be taken in designing porous-coated total joint implants with titanium alloy.

Another disadvantage of titanium alloy is its lower hardness (in comparison, for example, to

the cobalt alloys). An ambiguous term, hardness encompasses a number of mechanical

properties, but mostly measures the material’s resistance to elastic and plastic deformation.

Several standard tests for measuring hardness exist, most involving the forced indentation of

a fixed geometry indentor into the surface of a material. Hardness measurements are

determined from the geometry of the resulting indentation (for example, the depth or the

circumference). Hardness can be empirically related to other properties, such as yield

strength, but in general it is most useful in terms of comparison between materials.

Microhardness measurements, for example, in which a diamond-tipped, pyramid-

shapedindentor is pressed into the surface under a 10-g load, show titanium alloy to be about

15% “softer” than cast cobalt alloy. The decreased hardness of titanium alloy that must be

considered in total joint applications is due to its wear resistance.

Clinical observations have demonstrated significant scratching and wear of total hip femoral

heads made from titanium alloy. Measurements of the levels of titanium and aluminum in the

tissues and fluids taken from the hip joint have confirmed the release of significant amounts

of these elements from the femoral head. These observations suggest that titanium alloy that

has not undergone additional surface processing (for example, ion implantation) should not

be used as an articulating surface. Despite the longterm clinical evidence of the excellent

biocompatibility of titanium alloy, the concern that the release of cytotoxic elements such as

vanadium could cause local and systemic problems has led to the limited introduction of

other titanium alloys in which the vanadium has been replaced by more inert elements such

as niobium. Beta titanium alloys have also been advocated for orthopaedic implants. These

alloys have molybdenum concentrations greater than 10% to allow the beta phase to be stable

at room temperature. Beta alloys can be processed to possess lower elastic modulus (by about

20%) and slightly better crevice corrosion resistance than Ti-6Al-4V, while maintaining other

important mechanical properties at levels comparable or better than the conventional

aluminum-vanadium alloy. Together with excellent formability, the beta alloys are candidates

for a wide range of orthopaedic applications

Problems/test for with Biomaterials

Acute toxicity (cytotoxicity) arsenic

Sub chronic/chronic Pb

Sensitization Ni, Cu

Genotoxicity

Carcinogenicity

Reproductive &/or developmental Pb

Neurotoxicity

Immunotoxicity

Pyrogen, endotoxins