Calcium Phosphate Ceramics

Different phases of calcium phosphate ceramics are used depending upon whether a resorbable
or bioactive material is desired. These include dicalcium phosphate (CaHPO4) and
hydroxyapatite Ca10 (PO4)6(OH)2 [HA].

Applications include dental implants, skin treatments, gum treatment, jawbone reconstruction,
orthopedics, facial surgery, ear, nose and throat repair, and spinal surgery.

The mechanical behavior of calcium phosphate ceramics strongly influences their application
as implants. Tensile and compressive strength and fatigue resistance depend on the total
volume of porosity. Because HA implant have low reliability under tensile load, such calcium
phosphate bioceramics can only be used as powders, or as small, unloaded implants with
reinforcing metal posts, coatings on metal implants, low-bonded porous implants where bone
growth acts as a reinforcing phase and as the bioactive phase in a composite.

Calcium phosphate (CaP) biomaterials are available in various physical forms (particles or
blocks; dense or porous). One of their main characteristics is their porosity. The ideal pore size
for bioceramic is similar to that of spongy bone. Macroprosity (pore size >50μm) is
intentionally introduced into the material by adding volatile substances or porogens
(naphthalene, sugar, hydrogen peroxide, polymer beads, fibers, etc) before sintering at high
temperatures.

Microporosity is formed when the volatile materials are released. Microporosity is related to
pore size <10μm. Microporosity is the result of the sintering process, where the sintering
temperature and time are critical parameters. It has been demonstrated that microporosity
allows body fluid circulation whereas macroporosity provides a scaffold for bone colonization.
Average pore size diameter of 560μm is reported as the ideal macropore size for bone ingrowth
compared to a smaller size (300μm). The main difference between the different commercially
available BCP (biophasic calcium phosphate) are the microporosities, which are dependent on
sintering process.

Composites and Coatings

One of the primary restrictions on clinical use of bioceramics is the uncertain lifetime under
the complex stress states, slow crack growth, and cyclic fatigue that arise in many clinical
applications. Two solutions to these limitations are the use of bioactive ceramics as coatings or
in composites. Much of the rapid growth in the field of bioactive ceramics is due to
development of various composite and coating systems.

Composites have been composed of plastic, carbon, glass, or ceramic matrices reinforced with
various types of fibers, including carbon, SiC, stainless steel, HA, phosphate glass, and ZrO2.
In most cases the goal is to increase flexural strength and strain to failure and decrease elastic
modulus. The strongest composite achieved to date is A/W glass-ceramic containing a
dispersion of tetragonal Zirconia, which has a bend strength of 700MPa, and fracture toughness
of 4Mpam1/2.
Implant materials with similar mechanical properties should be the goal when bone is to be
replaced. Because of the anisotropic deformation and fracture characteristics of cortical bone,
which is itself a composite of compliant collagen fibrils and brittle HCA crystals, the Young's
modulus (E) varies between about 7 to 25GPa. The critical strain intensity increases from as

low as 600Jm-2 to as much as 5000J m-2 depending on orientation, age, and test conditions.

Most bioceramics are much stiffer than bone, many exhibit poor fracture toughness.
Consequently, one approach to achieve properties analogous to bone is to stiffen a compliant
biocompatible synthetic polymer, such as PE with a higher modulus ceramic second phase,
such as HA powders. The effect is to increase Young's modulus from1 to 8GPa and to decrease
the strain to failure from >90% to 3% as the volume fraction of HA increases to 0.5.

Thus, the mechanical properties of the PE-HA composite are close to or superior to those of
bone.

Another promising approach toward achieving high toughness, ductility and Young's modulus
matching that of bone was developed. This composite uses sintered 316 stainless steel of 50-,
100-, 200-μm or Titanium fibers, which provide an interconnected fibrous matrix which then
impregnated with molten 45S5 bioglass. After the composite is cooled and annealed, very
strong and tough material results, with metal to glass volume ratio between 4/6 to 6/4.

Stress enhancement of up to 340MPa is obtained in bending with substantial ductility of up to

10% elongation, which bends 90o without fracturing.

Coatings

A biometric coating, which has reached a significant level of clinical application, is the use of
HA as a coating on porous metal surfaces for fixation of orthopedic prostheses. This
approach combines biological and bioactive fixation. Though a wide range of methods have
been used to apply the coating, plasma spray coating is usually preferred. The table below
lists the bioceramic coatings:
1- Carbon
The medical use of pyrolytic carbon coatings on metal substrates were used in heart surgery.
The first time the low-temperature isotropic (LTI) carbon coatings were used in humans was a
prosthetic heart valve.

Almost all commonly used prosthetic heart valves today have LTI carbon coatings for the
orifice and/or occluder because of their excellent resistance to blood clot formation and long
fatigue life. More than 600,000 lives have been prolonged through the use of these bioceramic-
in-heart valves.

Three types of Carbon are used in biomedical devices:

1- Low temperature Isotropic (LTI);

2- Ultra low Temperature Isotropic (ULTI);
3- Glassy Carbons.

The LIT, ULTI and glassy carbon are sub-crystalline forms and represent a lower degree of
crystal perfection. There is no order between the layers such as there is in graphite; therefore
the crystal structure of these carbons is two-dimensional. Such a structure, called turbostratic,
has densities between 1400-2100 kg.m-3.

High density LTI carbons are the strongest bulk forms of carbon and their strength can further
be increased by adding silicon.
ULTI carbon can also be produced with high densities and strength, but it is available only as
a thin coating (0.1 to 1μm) of pure carbon.

Glassy carbon is inherently a low density material and, as such, is weak. Its strength cannot be
increased through processing.

The turbostratic carbon materials have extremely good wear resistance, some of which can be
attributed to their toughness, i.e. their capacity to sustain large local elastic strains under
concentrated or point loading without galling or incurring surface damage.

Another unique characteristic of the turbostratic carbons is that they do not fatigue. The
ultimate strength of turbostratic carbon, as opposed to metals, does not degrade with cyclical
loading.

Carbon surfaces are not only thromboresistant, but also appear to be compatible with the
cellular elements of blood'; they do not influence plasma proteins or alter the activity of plasma
enzymes. One of the proposed explanations for the blood compatibility of these materials is
that they absorb blood proteins on their surface without altering them.

2- Hydroxyapatite (HA)
A second bioceramic coating which has reached a significant level of clinical applications is
the use of HA as a coating on porous metal surfaces for fixation of orthopedic prostheses.

Resorption or biodegradation of calcium phosphate ceramic is caused by:

(i) Physiochemical dissolution, which depends on the solubility product of the material
and local pH of its environment;

(ii) Physical disintegration into small particles due to preferential chemical attack of grain
boundaries;

(iii) Biological factors, such as phagocytosis, which causes a decrease in local pH.
The rate of biodegradation increases as:

(a) Surface area increases;

(b) Crystallinity increases;
(c) Crystal perfection decreases; (d) Crystal and grain size decrease;

(e) Ionic substitutions of CO2− , Mg2+ , Sr2+ in HA take place.

Factors which result in a decreasing rate of biodegradation include:

(1) F-substitution in HA;

(2) Mg2+ substitution in β-TCP [β −Ca3(PO4)2 , β-tricalcium phosphate];

(3) Decreasing β-TCP/HA ratios in biphasic calcium phosphate.

Because of these variables, it is necessary to control the microstructure and phase state of
resorbable calcium phosphate bioceramic in addition to achieving precise compositional
control to produce a given rate of resorption in the body.

Natural Composites

Natural occurring composites are within us all. On the macroscale, soft and hard tissues are
formed from a complex structural array of organic fibers and matrix.

Soft tissues are formed from elastic (elastin) and non-elastic fibers (collagen) with a cellular
matrix between the fibers. Biological structures, such as tendon, linking muscles to bone,
are low in elastin, thus allowing muscle movement to be translated to the bone. However,
ligaments linking bone to bone are high in elastin allowing movement between bones but
retaining sufficient support to stop joints dislocating.
Hard Tissue

The two most important naturally occurring forms of bone required for structural stability are
termed cancellous (spongy bone) and cortical. Spongy bone is a sponge-like structure, which
approximates to an isotropic material. Cortical bone is highly anisotropic with reinforcing
structures along its loading axis and a highly organized blood supply called Haversian system.

All hard tissues are formed from the four basic phases, shown below. The relative fractions of
each phase vary between bone types and conditions. The relative percentages of mass for a
typical cortical bone are included.

The collagen fibers provide the framework and architecture of bone, with the HA particles
located between the fibers. The ground substance is formed from proteins, polysaccharides,
and musco-polysacharides, which acts as a cement, filling the spaces between collagen fibers
and HA mineral.

The microstructural level has two possible forms:

1- Woven bone is an immature version of the more mature lamellar form. Woven bone is
formed very rapidly and has no distinct structure.

2- Lamellar bone is formed into concentric rings called Osteons with central blood supply or
Haversian systems. Each osteon is formed from 4-20 rings, with each ring being 4-7 mm thick
and having a different fiber orientation. The arrangement of different fiber orientations in each
layer gives the osteon the appearance of successive light and dark layers. In the centre of the
rings, there is a Haversian canal which contains the blood supply. Whilst the outer layer is a
cement layer formed from ground substance, it is less mineralized than the rest of the bone and
has no collagen fibers. Consequently, the cement line is a site of weakness.

Synthetic Bone Grafting Materials

These materials must be:

1- Biocompatible with host tissues, i.e.

a- non-toxic;
b- non-allergic;
c- non-carcinogenic; d- non-inflammatory

2- Able to stimulate bone induction;

3- Resorbable following replacement by bone;
4- Radio-opaque;
5- Capable of withstanding sterilization
6- Inexpensive and stable to variation of temperature and humidity;

7- It has sufficient porosity to allow bone conduction and growth.