Ch11
Biomaterials
黃暉程 蔡昇翰 熊翌成 游琇婷
 Definition
of Biomaterial
 Classes of Materials
 Crystal
Structure
 Mechanical Behavior
 Wear
Resistance
 Calcium-phosphate
 Bioactive
Glasses
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P86994191
游琇婷
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
A biomaterial is a nonviable material
used in medical device, so its intended to
interact with a biological systems.
 Biomaterials
are manufactured substitutes
for natural tissues. They are used in
implants or catheters.artificial organs.
drug delivery. wound dressing.
4
 Biomaterials
can be conveniently grouped
into three classes: metals, polymers, and
ceramics.
 Composites, representing another group of
materials, consist of combinations of two or
more metals, polymers, or ceramics.
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 metallic
bond
 As the electrons can move easily in metals,
making metals easily deformable.
 The independent electrons in the metallic
bonds can quickly transfer electric charge
and thermal energy.
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 it
is often desirable to coat metallic
implants with a bioactive ceramic film in
order to improve implant fixation
 the difference in the thermal expansion
coefficient between metals and ceramics
results in interfacial shear stresses that can
create microcracks at the interface during
cooling subsequent to plasma-spraying a
calcium phosphate coating onto a metallic
implant.
http://www.gordonengland.co.uk/img/ps1.gif
http://www.azom.com/work/amvxiMoDGM6y1N7EBNXZ_files/image003.jpg 7
 As
the mechanical properties (and also the
chemical and physical properties) of metals
can be improved by alloying, most metals
used in orthopedic surgery are alloyed.
Obvious examples are the alloys based
either on titanium or on cobalt.
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 Ceramics
are refractory polycrystalline
compounds
•
•
•
•
•
•
Usually inorganic
Highly inert
Hard and brittle
High compressive strength
Generally good electric and thermal insulators
Good aesthetic appearance
 Applications:
• orthopaedic implants
• dental applications
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Synthetic HA
Bone HA
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 covalent
bond
 These large molecules contain many
repeating units, from which comes the
word polymers.
 The polymeric molecular arrangements
can be linear, branched, or cross-linked
catheters
artificial trachea
application
properties and design
requirements
polymers used
dental
•stability and corrosion resistance, plasticity
•strength and fatigue resistance
•low allergenicity
PMMA-based resins for
fillings/prosthesis
polyamides
poly(Zn acrylates)
ophthalmic
•gel or film forming ability, hydrophilicity
•oxygen permeability
polyacrylamide gels
PHEMA and copolymers
orthopedic
•strength and resistance to mechanical restraints
and fatigue
•good integration with bones and muscles
PE, PMMA
PL, PG, PLG
cardiovascular
•fatigue resistance, lubricity, sterilizability
•lack of thrombus, emboli formation
•lack of chronic inflammatory response
silicones, Teflon,
poly(urethanes), PEO
drug delivery
•appropriate drug release profile
•compatibility with drug, biodegradability
PLG, EVA, silicones, HEMA,
PCPP-SA
sutures
•good tensile strength, strength retention
•flexibility, knot retention, low tissue drag
silk, catgut, PLG, PTMC-G
PP, nylon,PB-TE
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 Composite
materials are solids that
contain two or more distinct constituent
materials or phases
 Most composite biomaterials have been
developed to enhance mechanical and
biocompatibility behavior.
 The shape of the phases in a composite
material is classified into three categories.
platelet
fiber
particle
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 Some
applications of
composites in biomaterial
applications are:
 (1)
dental filling
composites
 (2) reinforced methyl
methacrylate bone
cement
 (3) orthopedic implants
with porous surfaces.
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熊翌成
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 單位晶胞
(unit cells)
− 結晶結構的最小重複實體
− 可代表晶體結構的對稱性晶胞
− 晶體結構的基本單元
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 Body-centered
cubic (BCC)
體心立方結晶構造
 Face-centered cubic (FCC)
面心立方結晶構造
 Hexagonal close-packed (HCP)
六方最密堆積結晶構造
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 鄰近原子數目(配位數)=8
 每個立方格子含有2個原子
 原子填充率=0.68
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 晶體結構中另外兩個重要的特性是配位數
(coordination number)和原子填充率 (atomic
packing factor, APF)。對金屬而言，每一原
子具有相同的最鄰近或接觸原子的數目，就
是配位數的定義。
 另外APF是單位晶胞中固態球體的體積分率，
假設單位晶胞具有原子硬球模型時，則APF
的定義為單位晶胞中原子的體積除已全部單
位晶胞的體積所得之因子。
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左圖中，a 為晶胞立方格子單位長度，
R為原子半徑，二者關係可用下列式
子表示，穿過體心的對角線為鐵原
子排列最緊密的方向。
4𝑅 = 3𝑎
𝑎=
4𝑅
3
考量原子填入晶胞所占有的空間，可將原子填充率以下列式子表示：
晶胞內所含的原子體積
原子填充率 =
晶胞體積
因此，體心立方晶體的
2 4𝜋𝑅3 3
2 4𝜋𝑅3 3
原子填充率 =
=
3 = 0.68
𝑎3
4𝑅 3
在後面不同的晶體結構比較，我們會發現晶體的原子填充率與其最鄰近的
原子數目有關。最鄰近原子數目為8的體心立方結構，不是原子最緊密堆
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積的結構。
 鄰近原子數目(配位數)=12
 每個立方格子含有4個原子
 原子填充率=0.74
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由左圖可見，通過面心的格子
對角線為緊密排列方向。若晶
胞立方格子單位長度為a，原子
半徑為R，其間關係可以下式表
示：
4𝑅 = 2𝑎
𝑎=
4𝑅
2
2 4𝜋𝑅3 3
2 4𝜋𝑅3 3
原子填充率 =
=
= 0.74
3
3
𝑎
4𝑅 2
面心立方晶體的原子填充率(0.74)較體心立方晶體者(0.68)為高。事
實上，具有最鄰近原子數為12的金屬晶體，其原子排列為空間最緊
密的一種堆積結構。另一種最緊密堆積結構，則為六方最密堆積結
構。
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不是所有金屬的單位晶胞都具有立方對稱，第三種常見的金屬晶體結
構是具有六方立體晶格的單位晶胞，稱之為六方緊密堆積(hexagonal
close-packed 簡稱HCP)。在每一單位晶胞中包含有6個原子，計算方式
為每個單位晶包含有12個頂面和底面角落原子，其中每一個原子的六
分之一包含在這個單位晶胞中，另外晶胞亦包含2個中心平面原子的
每一個的二分之一和所有3 個中間平面的內部原子。
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• 鄰近原子數目(配位數)=12
• 每個立方格子含有6個原子
• 原子填充率=0.74
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
利用x, y, z座標系統的六個參數定義單位晶胞：
• 單位晶胞各邊以x、y和z座
標軸表示，各邊邊長以a、
b、c表示，邊與邊之間的
夾角以α、β和r表示。
• 此六個參數為三個邊長a、
b和c，及三個軸的夾角α、
β和γ，這些亦稱為晶格參
數(lattice parameters)。
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 All
metals, most ceramics, and some polymers
crystallize when they solidify.
 A crystalline material is characterized by longrange order and an infinitely repeating unit cell
of atoms/ions.
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TABLE 11-2. Some Materials and Their Representative Equilibrium Crystal
Structure at Body Temperature
Material
Structure
Cobalt-chromium alloy
FCC
Stainless steel AISI 316L
FCC
Titanium (Ti)
HCP
Tantalum (Ta)
BCC
Niobium (Nb)
BCC
Gold (Au)
FCC
Alumina (Al2O3)
HCP
Hydroxyapatite [Ca10(PO4)6(OH)2]
HCP
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 Hydroxyapatite
(HA), a bone bioactive ceramic.
 HA structure is considerably complex, as a
result of which, displacement of atoms within
the lattice is difficult.
 Thus the structure is resistant to deformation
and, when overloaded, fractures rather than
deforming permanently.
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34
 As
the extent of polymerization increases and
the molecular chains become longer, the
relative mobility of the chains in the structure
decreases.
 As a result, alignment of the chains and
formation of long-range order is difficult.
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36
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 Factors
affecting the strength of polymers
further include chemical composition, side
groups, cross-linking, copolymerization, and
blending.
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 Mechanical
properties including elasticity and
strength are important properties to consider
in the selection of a material for a specific
implant design.
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TABLE 11-3. Typical Values of Tensile Strength for Various Materials at
Room Temperature
Material
Diamond
Tensile Strength (MPa)
1.05 × 106
Kevlar
4,000
High-strength carbon fiber
4,500
High-tensile steel
2,000
Superalloy
1,300
Spider webs (drag line)
1,000
Ti-6Al-4V
860
CoCr alloy (F75)
655
Aluminum
570
Titanium (grade 4)
550
316L SS (F745, annealed)
485
(Cold forged)
1,351
Bone
200
Nylon
100
Rubber
100
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TABLE 11-5. Elastic Properties of Some Typical Materials
Modulus of Elasticity
(MPa)
Materials
Directionality Properties
Cortical bone
Anisotropic
Longitudinal axis, 17,000
Trabecular bone
Anisotropic
Longitudinal axis of femur
intertrochanteric 316 ± 293
High-density polyethylene
Isotropic
410-1,240
PMMA
Isotropic
3,000-10,000
Stainless steel AISI 316L
Isotropic
200,000
Cobalt-chromium alloy
Isotropic
220,000
Titanium
Isotropic
107,000
Ti-6A1-4V
Isotropic
110,000
Carbon fiber-reinforced graphite
fibers
Anisotropic
Parallel to unidirectional 140,000
Single crystal, 362,700
Alumina
Isotropic
Polycrystal, 408,900
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 簡單破壞是指一個物體在低溫下(相對於熔
點)，受到施加靜態應力(即應力為常數或隨
時間緩慢改變)，分裂為兩個或更多的碎片。
 對工程材料而言，依據材料發生塑性變形的
能力將其分類，有兩種可能的破壞模式：延
性(ductile)和脆性(brittle)。延性材料在破壞
之前通常出現高能量吸收的大量塑性變形，
而脆性材料的破壞幾乎沒有塑性變形，只有
低能量吸收。
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 破斷面上大量的塑性變形，就是延性破壞的證據。
受到拉伸時，高度延性金屬破斷面會頸縮至一點。
 延性材料的裂紋是穩定的(沒有增加外在應力即不
會生長)，由於不是突然及災難性的破壞，所以這
種破壞模式較能接受。
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 脆性破壞藉著快速的裂紋生長，在幾乎沒有
變形的情況下就發生了。裂紋的運動方向幾
乎是垂直於施加的拉伸應力，產生出一個相
當平坦的破斷面。
45
(a) 鋁的延性破壞，(b) 中碳鋼的脆性破斷。
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對多數脆性結晶材料而言，裂紋成長相當於沿著特定結晶平面，相
繼重覆地打斷原子鍵，這個過程稱為劈裂(cleavage)，這種形式的破
壞稱為穿晶破壞(transgranular；或稱 transcrystalline)，因為破壞
裂縫穿越晶粒而成。
(a) 穿晶破壞時，裂紋沿晶粒內部前進的剖面圖示。
(b) 延性鑄鐵的掃描電子破斷面照片顯示穿晶破斷面。
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有些合金的裂紋沿晶界前進；此形式的破壞稱之為沿晶破壞
(intergranular)。
(a) 沿晶破壞時，裂紋沿晶界前進的剖面圖示。
(b) 掃描電子破斷面照片顯示一沿晶破斷面。放大 50 倍。
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應力集中 Stress Concentration
 材料在正常情況下，其表面或內部總是存在
非常微小的瑕疵或裂縫，這些瑕疵對於破壞
強度是一種損傷，因為施加應力會放大或集
中於裂紋的尖端，應力放大的量取決於裂紋
的方向和幾何形狀。這些瑕疵由於它們所在
之處有放大應力的能力，因此有時稱為應力
集中源(stress raiser)。
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 在室溫下，結晶和非結晶陶瓷在受到拉伸負荷
時，幾乎在塑性變形發生之前就已破壞。
 當應力基本上是靜態的，陶瓷材料的破壞是藉
著裂紋的緩慢前進來發生的，這種現象稱為靜
力疲勞(static fatigue)或是延遲破壞(delayed
fracture)。
 同一種脆性陶瓷材料做的不同試片所測得的破
壞強度值常有所變動。對於壓應力而言，就沒
有因為瑕疵造成的應力放大現象，因此，脆性
陶瓷的抗壓強度比抗拉強度高得多(相差 10 倍
的等級)，所以常被用來承受壓負荷。
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 Fatigue
cyclic load
fatigue strength
fatigue limit
 靜力疲勞(Static
fatigue)
或稱延遲破壞(delayed fracture)
通常發生在陶瓷材料上
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TABLE 11-4. Fatigue Properties of Implant Metals
ASTM Designation
Stainless
steel
F745
Annealed
221-280
F55, F56, F138, F139
Annealed
241-276
30% Cold worked
310-448
Cold forged
820
As-cast/annealed
207-310
P/M HIPa
725-950
F799
Hot forged
600-896
F90
Annealed
Not available
44% Cold worked
586
Hot forged
500
Cold worked, aged
689-793 (axial tension R = 0.05,
30 Hz)
F67
30% Cold-worked
grade
300
F136
Forged annealed
620
Forged, heat treated
620-689
Co-Cr alloys
F75
F562
Ti alloys
a
Condition
Fatigue Endurance Limit
(at 107 Cycles, R = -1) (MPa)
Material
P/M HIP, Powder metallurgy produced, hot-isostatically pressed.
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黃暉程
 Ultrahigh-molecular-weight
Polyethylene
(UH-MWPE) (MW above 2 × 106 g/mol)
 The
success of UHM-WPE is due to its
favorable properties, including abrasion
resistance, impact strength, low coefficient
of friction, chemical inertness, and
resistance to stress cracking.
 It
has long been used for total hip
prostheses and knee prostheses.
 In
the majority of contemporary total joint
replacements, a metallic component articulates
against UHMWPE.
 As
a patient may be expected to take an
average of 1,200 steps per day, the joint
replacement is expected to withstand millions
of loading cycles during its service lifetime.
 In
total hip replacements, where the
articulating geometries consist of conforming
spherical surfaces, the wear occurs at a
microscopic length scale (µm or less).
 Wear
resistance has been related to its
resistance to multidirectional stresses.
 In
total knee replacements, where the
articulating surfaces consist of nonconforming
cylindrical, toroidal, or flat surfaces, the wear
process includes various surface damage
mechanisms ranging from pitting and
delamination to burnishing and adhesive wear.
 Adhesive
wear is a process in which surface
asperities of the polymer adhere to the metal
surface and subsequently are torn off.
 As
a result, either a polymer film is formed on
the metal surface or polymer particles are
released and entrapped in the joint.

 Such
particles as well as remaining cement
fragments can generate abrasive wear, a
second mode of wear.
 Abrasive
wear can also be generated by
asperities on the gliding metal partner.
A
third mechanism is fatigue wear; as a result of
creep or plastic flow, folds or cracks are formed
that cause small polymer particles to break off.
 It
is generally accepted that particulate debris
generated by mechanical wear of prosthetic
components stimulates the generation of a
pseudosynovial membrane at the interface
between implant and bone and the infiltration
of fibrocytes and macrophages.
 In
the presence of debris, these cells release
various cytokines and mediators (such as IL-1β,
TNFα, collagenase, and prostaglandin E2)
 These
cytokines have been shown to be
involved in bone resorption by activating
osteoclasts.
 A recent study showed a positive correlation
between cytokine concentration in the
loosening membrane and the degree of
underlying osteolysis.
 Many reports have investigated the failed total
joint arthroplasty and demonstrated that
phagocytosis and cell mortality increase with
particle size and concentration.
 The
mechanism that results in this cell death
remains unknown, although a potential role
for apoptosis in the pathogenesis of wear
debris-associated osteolysis has recently
been suggested.
 Aseptic
loosening, which is the single most
common cause for long-term failure of TJA, is
associated with periprosthetic osteolysis with
the incidence of up to 25% of implant
recipients.
 Multiple
factors can affect polyethylene wear
and the production of wear debris in vivo after
joint arthroplasty.
 Such
factors include the roughness and
material of the femoral head, the method of
polyethylene sterilization, and the mechanical
properties of the polyethylene itself.
 Recent
laboratory studies have confirmed that
the wear resistance of UHMWPE can be
significantly increased when applying
additives or high-dose irradiation.
 Many
studies reporting extremely low
quantities of wear, typically as low as 2.0 mm3
per million cycles for highly cross-linked
polyethylenes.
 High-density, high-purity
alumina is used in
load-bearing hip prostheses because of its
outstanding wear resistance and excellent
corrosion resistance.
 It has a high Young's modulus and a hardness
second only to that of diamond.
 These properties have made the alumina-onalumina couple for femoral heads and
acetabular cups a materials combination of
considerable importance.
 Most
alumina devices are very fine-grained
polycrystalline α-Al2O3 produced by pressing
and sintering at high temperature (1600°C to
1700°C).
 A very small amount of MgO (<0.5%) is used to
aid sintering and limit grain growth during
sintering.
 Alumina with an average grain size below 4 µm
and a purity greater than 99.7% exhibits
excellent flexural and compressive strength.
 Studies
have reported excellent performance
of the alumina-on-alumina bearing in terms of
low annual wear (<5 µm).
 The long-term friction of an alumina-alumina
joint prosthesis decreases with time and
approaches the value of a normal joint.
 This lead to wear on alumina articulating
surfaces being nearly 10 times lower than on
polyethylene surfaces gliding against metallic
heads.
 The
pseudosynovial tissue obtained from
around retrieved uncemented ceramicceramic prostheses have identified numerous
alumina ceramic particles with a mean size of
5 µm.
 Henssge
et al. observed particles up to 5 µm
in diameter in the periprosthetic tissues from
around cemented alumina-alumina
prostheses.
 The
bimodal size range of alumina ceramic
wear debris overlapped with the size ranges
commonly observed with metal particles (10 to
30 nm) and particles of UHMWPE (0.1 to 1000
µm).
 It
is possible that the two types of ceramic wear
debris are generated by two different wear
mechanisms in vivo.
 Under
normal articulating conditions, relief
polishing wear and very small wear debris is
produced, while under conditions of
microseparation of the head and cup and rim
contact, intergranular and intragranular
fracture and larger wear particles are
generated.
 Petit
et al. compared the macrophage response
with identically sized particles of alumina
ceramic (Al2O3) and UHMWPE in terms of TNFα release and induction of apoptosis of J774
mouse macrophages.
 The
stimulation of TNF-α release was much
greater (8 to 10 times higher) with UHMWPE
than with Al2O3.
 It
could be possible that the ability of Al2O3
particles to induce macrophage apoptosis
may explain the lower TNF-α release
observed with these particles and explain the
differences seen in osteolysis patterns of
ceramic-ceramic versus metal-PE
articulations.
 The
induction of macrophage apoptosis may
therefore be a desirable therapeutic
endpoint.
 The
human body is a very aggressive
environment for metals.
 Corrosion
resistance of a metallic implant
material is an important aspect of its
biocompatibility.
 Metallic
biomaterials are normally considered
to be highly corrosion resistant because of the
presence of an extremely thin passive oxide
film that spontaneously forms on their surfaces.
 The
properties of these passive oxide films
depend to a large extent on their structure and
chemistry.
 There
are two Co-Cr alloys extensively used
in artificial joints for heavily loaded joints
such as knees and hips: the castable Co-CrMo alloy and the Co-Ni-Cr-Mo alloy, which is
usually wrought by (hot) forging.
 The inhomogeneous microstructure of the
cast Co-Cr-Mo alloy renders it more
susceptible to corrosion than the forged alloy.
 The
chromium is a reactive element and is
added to produce a stable firmly adherent
protective chromium oxide surface layer. It also
enhances the solid strengthening of the alloy.
 The
molybdenum is added to produce finer
grains, which results in higher strengths after
casting or forging. It also enhances the
corrosion resistance of Co-Cro alloys.
 The
metallic products released from the
prosthesis because of wear and corrosion
may impair organs and local tissues, and
moreover, some alloys with certain amount of
Co can be toxic in the body.
 Low
wear has been recognized as an
advantage of metal-on-metal hip articulations
because of their hardness and toughness.
 Both
c.p. Ti and Ti-6Al-4V possess excellent
corrosion resistance for a full range of
titanium oxide states and pH levels.
 However, it
derives its resistance to corrosion
by the formation of a solid oxide layer to a
depth of 10 nm.
 Under
in vivo conditions the oxide, TiO2, is a
very stable reaction product.
 Stainless
steel contains enough chromium to
confer corrosion resistance by passivity. The
relatively resistant varieties of stainless steel
are the austinitic types 316, 316L, and 317,
which contain molybdenum (2.5% to 3.5%).

 The
corrosion resistance can be enhanced by
increasing the thickness of the protective
oxide using concentrated nitric acid
(“passivation”), by boiling in distilled water,
or by electrochemical means (anodization).
 Corrosion
of an implant in the clinical setting
can result in symptoms such as local pain and
swelling in the region of the implant, with no
evidence of infection.
 Wounds
and infections can significantly
change pH.
 Different
parts of the body undergo
different types and rates of corrosion.
 Bioactive
materials including calcium
phosphate ceramics and bioactive glasses
bond to bone and enhance bone tissue
formation.
 The
forms of calcium phosphate ceramics
most widely used are tricalcium phosphate
[Ca3(PO4)2, whitlockite], tetracalcium
phosphate (Ca4P2O9), and hydroxyapatite
[Ca10(PO4)6(OH)2] (HA).
 The
variation of mechanical properties is the
result of variation of density and crystalline
structure.
 Sintering of calcium phosphate ceramics is
usually carried out in the range of 1,000°C to
1,500°C.
 Depending on the final firing conditions, the
calcium phosphate can be hydroxyapatite or βwhitlockite.
 The
phases formed at high temperature
depend not only on the sintering temperature
but also on the partial pressure of water in the
sintering atmosphere.
 The
temperature range of stability of HA
increases with the partial pressure of water,
as does the rate of phase transition of
tricalcium phosphate or tetracalcium
phosphate to HA.
 Schepers
et al.
(1991): granules of
HA implanted in
bone tissues of the
beagle mandible.
After 3 months, bone
tissue has grown
along the particle
surface, but only
over a distance of
about 1 mm.
 From
6 months on,
the particle surfaces
have a moth-eaten
appearance caused
by multinucleated
cell resorption.
 This
experiment showed beyond doubt
that dense, stoichiometric HA is
osteoconductive in vivo.
 However, it
also revealed that this
material displays only a limited bone
bioactivity.
P86991177
蔡昇翰
 Direct
implant-bone bonding, better fixation,
longevity
ingrowth with HA lining:
pronounced at 2 ~ 4 weeks, faster
rehabilitation
 ↑bone
 Different
ceramic characteristics  various
bonding strength
CAP1: hydroxyapatite
powder + CPC:
poly(lactic acid)
CAP2: hydroxyapatite
powder
CAP3:
oxyhydroxyapatite
Enhanced bone tissue growth fixation
Effect differed among coatings
Ducheyne P, Biomaterials 1990







D'Antonio: 316 hips, early pain relief, rapid restore function
Hernandez: 52 hip arthroplasties, 11-year survival rate 92.3%
Geesink: 118 arthroplasties, 99-100% survival rate at 6 years
Vidalain: 0.97 survival rate of femoral stem at 10 years
Excellent function result (Harris hip scores), particularly pain
63% totally pain free, recovered normal motion & function
Conclusion: osteoconductive coatings have excellent
performance
 HA: toxic
(-), inflammatory or allergic reactions
(-), carcinogenic (-)
1.
2.
3.
4.
HA resorption & long-term stability
Osteolysis
Polyethylene wear & potential risk of
granuloma
Difficulty of extracting an HA-coated stem

HA: considered insoluble at neutral pH, not be degraded

Aebli et al: in proximally femoral, complete coating degraded
after 9.5 years, no interposing fibrous tissue

Resorption: both chemical dissolution & cell-mediated
degradation (not all well known)

Loss of HA coating: no negative effect on osseointegration

Conclusion: degradation not adversely affect long-term
fixation

HA decomposed to several phase (α-/β-TCP, CaO, TTCP,
oxyapatite, amorphous calcium phosphate…) with plasma

Factor: plasma parameter, coating environment
(dielectric constant, pH), cell activity (osteoclast,
lysosome)…

Coating degradation  calcium-phosphate particle stacking



Bloebaum: histologic exam  HA particles,
inflammatory reactions, & osteolysis
Calcium particle migration  trigger third-body
water  inflammation, osteolysis
Resorption  void between bone matrix 
mechanical instability

Reikeras et al:
• 155 pt, 39 cups revised (mechanical loosening)
• 9 radiolucent lines, 2 focal osteolysis, none had symptoms

Still Good overall outcome of Ca-P coatings

Next-generation biomimetic HA coatings: 3-D ingrowth
structures, combinations with antibiotics, growth factors…
Silica (SiO2) based glasses: network former
 Alkali metals (Na, K) or earth metals (Ca, Mg):
modifier
 Ratio  solubility, bioactivity & resorbability



Lower mechanical strength (amorphous
structure): unsuitable for load-bearing
applications
• Low modulus of elasticity of 30~35 Gpa: close
to cortical bone (7~25 GPa)
Form a strong interfacial bond with adjacent
tissues
• Bone-biomaterials interface developed in vivo
• Immersion in simulated physiological fluids or
cell-containing media in vitro
-Si-O-Na + H+ + OH → -Si-O-H + Na+ (solution) + OH -Si-O-Si- + H2O →-Si-OH + HO-Si O-Si-OH + HO-Si-O →O-Si-O-Si-O + H2O
 Ca2+ & PO43- migrate to surface through SiO2-rich
layer  forming CaO-P2O5-rich layer on top of
SiO2 gel layer

Absence of proteins & bone cells: amorphous Ca-P
layer crystallizes into apatite on surface
 + protein  reaction not occur in vivo or in vitro
 + osteoblasts  HA surface form again (cellular
activity)

 HA
develop in acellular SBF // ability to bond
to bone
 Bioactivity
of biomaterials is evaluated based
on ability of HA development in vitro
 But…tissue
fluid ≠ SBF
 And different parametric conditions in vivo & in
vitro
1.
2.
3.
4.
5.
6.
7.
8.
9.
10.
11.
Dissolution from the ceramic
Precipitation from solution onto the ceramic
ion exchange and structural rearrangement at the
ceramic-tissue interface
interdiffusion from the surface boundary layer into
the ceramic
solution-mediated effects on cellular activity
deposition of either the mineral phase or the organic
phase, without integration into the ceramic surface
deposition with integration into the ceramic
chemotaxis to the ceramic surface
cell attachment and proliferation
cell differentiation
extracellular matrix formation
 Bone
bonding & bone tissue ingrowth
enhancement: multiple, parallel, & sequential
reactions at material-tissue interface
 Hydroxyapatite surface: lead to biologically
equivalent apatitic surface on implanted
material
 From material: solution-mediated & surfacecontrolled effect on cellular activity, organic
matrix deposition, & mineralization
 All
lead to gradual incorporation of bioactive
implant into developing bone tissue
& biochemical property: glassceramic +/- bioactive glass
 Thermal treatment:
• 550°C ~ 680°C  Na2Ca2Si3O9 crystals
• ~ 800°C  calcium phosphate crystals
similar to HA
 ↑mechanical
 In
MgO-CaO-SiO2-P2O5 
apatite/wollastonite (A-W) glass ceramic
• High mechanical strength: for load-bearing
prostheses
• ↓rate of bone bonding to material
 Not
homogeneous, immiscible glassy phases
of different chemical constituents
 ↓serum
protein adsorption onto material
surface
Slow rate of bone bonding to bioactive glass
ceramic
 Delayed
formation of HA surface layer
necessary for bone bonding
•
•
•
•
Not heal fractures, major bone loss, & bone
tumors
Porous scaffold seeds cells for tissue regeneration
Patient's own cells (not immunosuppression)
Scaffolding material:
– Biocompatible & biodegradable (nontoxic &
easily excreted by metabolic pathways)
– Strong mechanics to maintain structural integrity
during culture
– Easy to fabricate into a desired shape & porous
architecture
– Osteoconductive
1.
Poly(α-hydroxyl acids):
• Degradation products  ↓pH  ↑polyesters'
degradation rates & inflammation
• Weak mechanics: limits for bone
regeneration
• Incorporation with HA: compact, more stable
pH
2.
HA ceramic: suitable for bone tissue scaffold,
but resorbs very slowly
3.
Bioactive glass:
– Stimulates osteoprogenitor cell function &
possesses controlled resorbability
– Ca-P layer: selective fibronectin adsorption
 ↑osteoblast adhesion & activity
•
Dual layer (Ca-P + serum protein)  abundant
& expeditious bone tissue
•
•
•
Osteoblast adherence to biomaterial surface:
– Topography, chemistry, surface energy
– Tripeptide (arginine-glycine-aspartic acid -RGD): located in cell-binding domain of
many adhesion molecules
Reorganization of cytoskeletal proteins 
flattening & spreading of cell  regulate cell
behavior
Limit cell attachment period: protein
desorption, exchange & denaturation,
proteolysis, difficulties in controlling sequence
•
•
Covalent immobilization of active peptide is
necessary
Chemistry of surface layer: can modify peptide
conformation & its interaction with cells
• RGD motif grafted to quartz surface
• Different RGD & FHRRIKA binding domain ratio
• Oxide surface activated by aminopropyl
triethoxysilane
•
Immobilized peptides with RGD: ↑ cell
attachment to polymeric & ceramic surfaces