Bioceramics for hard tissue replacements

Clinical Materials1987;%2: 181-206

its

for hard tissue replacement

~~~~~ar~ Institute of Clinrcai Materials, Osaka

No materials placed within a living tissue can be considered to be completely inert, However, stable ceramics do not suffer from corrosion or degeneration as do metals or plastrcs. Extensive progress in ceramic technology has developed new functional bioceramics with a high degree of biocompatibility or bioinertness, or with bioadhesiveness with bioactive and biodegradable characteristics which can accelerate new bone growth. However, ceramics with a characteristically high Young’s modulus are apt to produce biomechanical unbalance at the implant-tissue interface and cause chronic mechanical irritation which introduces tissue inflammation and probably conceals neoplastigenicity. To prevent such mechanical irritation, ceramic implants should be fixed strongly into bone tissue. Porous ceramics are effective in giving reliable fixation and have produced many successful oral and orthopaedic implants to date.

Introduction

Nonbiological materials have become indispensable in the management of hard tissue defects. As a result, in dentistry, a rather unique scientific field has developed which faces issues such as the choice of materials, the design of ap’pliances, the moulding of artificial materials and matching them to living tissues. Such studies have made numerous contributions to the development of biomaterials for use in the orthopaedic fields. Systematic biological evaluation of dental and surgical materials has been carried out by the author since 1965l At present, it is impossible to Address for correspondence: Professor H Kawahara, Institute of Clinical Materials, Higashiyama 4-8, Higashi Osaka 579, Japan.

make artificial materials which equate wholly with living tissues. Attempts to create implantable biomaterials of equal characteristics as those of the recipient’s tissues are likely to be in vain, but better substitutes must continue to be sought? Implant materials have to be:

1) noncytotoxic,

nonirritating, nonallergenic and noncarcinogenic; 2) biomechanically matched with the physical properties of the tissue; and 3) bioadhesive: implant materials which are not bioadhesive do not allow close attachment of the tissues surrounding the implant and may easily loosen.

182

H Kawahara

Bioinert and bioadhesive

materials

Materials closely similar to living tissue will be easily dissolved and digested in the host tissue, therefore implant materials with molecular structures similar to proteins or polysaccharides are unlikely to be stable. For example, nylon (polyester amide) is likely to become an allergen and is therefore easily digested and absorbed within the body (Figure l).” On the other hand, silicone rubber, polyethylene and polytetrafluoroethylene (Teflon), which have molecular structures completely different from biological substances, are generally more stable in the body. However, these polymers are hydrophobic and offer little adhesion to living cells and tissues in spite of their high stability. It is difficult to fix them firmly in the bone tissue and to protect the surrounding tissues from infection and to prevent epithelial downgrowth at the junction between dental implant CH 13

CH3 I CH3-

ii-

Cji3 0 -

Si-

0 -

CH3

Si-

I

I

CH3

CH3

I

and the gingival mucosa.3,4 They are not, therefore, suitable materials for use in the replacement of dental roots, or bone and joints. Indeed, there is no polymer which has both good adhesion to living cells combined with high stability in tissue fluids. However, this problem might be overcome. Recent research on antithrombonic polymers has shown that implantable plastics, with high stability, can be induced to adhere strongly to the surrounding tissues by the grafting of hydrophilic polymers onto the surface of the hydrophobic plastics. Inorganic materials, such as aluminium, indium, tin, silicon and chromium are comparatively stable in living tissue and exhibit almost no tissue irritability and have favourable adhesion to cells and tissues. 5 However, aluminium, indium , silicon and chromium are corroded easily and ionized if the pH at the site of the implant becomes acidic, as occurs in infection. Also, these metals are not

. . . . . . . . . I silicones

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Figure 1 structure

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Materials which resemble proteins and polysaccharides are more unstable in living tissues

in molecular

Bioceramics for hard tissue replacements

suitable for use in their pure form because of their poor mechanical strength. In contrast, titanium, zirconium and tantalum are more biomechanically compatible and fairly stable in living tissue.5,6 Tissue irritability and cytotoxic properties of metals do not necessarily coincide with their relative ionization. This may be because cellular responses differ depending upon the passive state of the (ceramic-like) oxidized film on the metal surface7 These oxidized films protect metals from corrosion in living tissue and produce a high degree af adhesiveness to the surrounding tissues8zg This also applies to the alloys of cobalt-chromiummolybdenum and iron-chromium-nickel. When 20% of chromium is mixed with cobalt and/or nickel, the cytotoxicity of these elements is masked, because the alloy surface is covered with a homogeneous oxidized chromium film in a passive state. 10,11 Titanium, zirconium, tantalum, cobalt-chrome alloys and iron-nickel-chromium-molybdenum alloys are widely used in the construction of surgical appliances and dental implants. In this titaniummagnesium-aluminium, respect, aIumi~i~m-va~adium,12 titanium-molybdenum,13 and titanium-nickel14 titanium-palladium, titanium-~ircQ~iurn alloy9 all deserve to be evaluated as they are light, stable, highly corrosion-resistant and have shape memory qualities, but the possible inflammatory and toxic effect of metallic implants is still of concern. It has been reported that elements of the metallic implants titanium, cobalt-chromium-molybdenum and iron-nickel-chromium have been known to dissolve into the surrounding tissue.16 In particular, shape memory alloys of titanium-nickel demonstrate nickel dissolution into living tissue during long term implantation.r7Js

ility of materials No material placed within a living tissue can be

considered to be completely inert. However, ceramics, by their very nature, do not suffer from corrosion as do metals or plastics. Extensive progress in ceramic technology has led to the discovery of a variety of materials whose chemical, physical and mechanical properties have a high

183

degree of stability and bioadhesiveness. 19,20This makes them suitable candidates for lon~te~m implants within living tissue. As mentioned above, the more the implant material resembles the host tissue in molecular structure, the more severe is the reaction in the living tissueZ1 (Figure 2). The reason for this is because of the difficulty the host experiences in discriminating immunologically between implant material and host tissue. Conversely, removal of the antigenic stimulus from implanted material or transplants results in a more inert nonbiological agent that the host easily recognizes as foreign. Ceramics which have molecular structures completely different from those of living substances are generally more stable inside the living body and have a high adhesion to the surrounding tissue as shown by in vitro and in vivo experiments.

/ I

METALS

I

PLASTICS

/

BIOLOGICAL

MATERIALS

Figure 2 The greater the difference between a material and the living substance, the more inert is the reaction within living tissue.

It is interesting to note that those materials which are more easily fabricated are often the most unstable, degradable, irritable and toxic in living tissue (Table 1). This is a function of material bonding, atomic structure and defect concentrations. In vitro stability decreases from selfcuring plastics to injection mouldable polymers. Metallic materials with higher melting points are more stable and inert in living tissue and they are more difficult to form. Ceramics with high sintering temperatures, e.g. alumina, zirconia,z2J3 silicon

184

H Kawahara

nitride and vitreous carbon have high stability in living tissue (Table 2), although vitreous carbon implants have not been used in clinical cases due to inappropriate designs and the aesthetic problem of its black colouration. Some bioinert ceramics possessing a relatively low crystallization and sintering temperature demonstrate decreased stability and mechanical strength after the implantation, compared to those bioactive ceramics possessing higher processing temperatures which are more stable in living tissue. From these findings, one may conclude that single crystal alumina of sapphire should have the highest stability and thus should be considered as suitable biomaterials for hard tissue replacement .2k26 Bioinert ceramics vs bioactive ceramics

Bioinert ceramics The stable ceramics such as alumina, zirconia, silicon nitride, vitreous carbon, and LTI carbonz72g exhibit high biocompatibility and adhesion to tissues and have a higher mechanical strength than bioactive ceramics. Complex designs for hard tissue implants are possible, because no change of mechanical properties or chemical components is observed even after long implantation times. In recent years, endosteal dental implants of sapphire, BioceramR (Kyocera, Japan)24 (Figure 3) and zirconia have been considered to be of particular interest for such applications. Bioactive ceramics Hench and Paschal130 have developed a glass with soluble additives called Bioglass (Si02 3045%) ONazO 24.5%) CaO 12.25-24.5%, P,05 6%) CaF, 6-12.25%). According to their report, insertion of this material into bone resulted in silica gel formation around the Bioglass and promoted collagen formation due to the dissolution of calcium and phosphorus from the bioglass surface into the adjacent tissue. The resultant chemical bond between the silica and collagen protein then enabled the crystallization of hydroxyapatite. Br8mer, Deutscher, Blencke and Strunz31 also have developed a new glass ceramic called CeravitarR (SiOZ 40-45%, CaO 20-30%) which contains Na,O 15-30%) P,O, 5-lo%),

K,O and MgO (S-10%) in order to match the ionic constitution of blood and body fluid. Normal bone formation and a tight adhesion between the new bone tissue and the glass ceramics without soft tissue encapsulation were observed. Mouroe et al.32 proposed the use of sintered, fine-grained, synthetic hydroxyapatite for dental implants, since this material resembles tooth substance. Fischer-Brandies and Dierlart33 implanted a dense hydroxyapatite (Ca,&Po,),(OH),) on unreamed canine cortical bone and concluded that a direct contact formed between the hydroxyapatite and bone without any connective tissue layer. However, these bioactive ceramics have shortcomings in their bending strength and impact strength (Table 2). CeravitarR has a strength of 98MPa and Bioglass probably less than 100 MPa. Fischer-Brandies and Dierlert reported that sintered dense hydroxyapatite ceramics have bending strengths of between 113 and 196 MPa. The latter value was, however, obtained from a small, highpolished sample. Therefore, it’s practical value would in fact be reduced to around 115 MPa. Kokubo et aZ.34reported that a new type of glass ceramic containing apatite and wollastonite (CaO 44.9%, SiOZ 34.2%, P,O, 16.3%) MgO 4.6%, CaF 0.5%). It has a consistently high bending strength of 157 + 8 MPa. Gross and Strunz35 presented their histological observations of the between glass suitable interfacial reaction ceramics and new bone, but the reported bending strength was still only 113-157. Strength is clearly a problem with the bioactive ceramics. Furthermore, the solubility of the bioactive elements may reduce the mechanical properties of longterm implants. It is believed that bioactive ceramics might form an organic bond with new bone following osteogenesis and make close adhesion or osteoankylosis between the new bone and the bioactive ceramics surface. These expected phenomena do not always occur, because osteogenesis is an extremely complex interfacial phenomenon.36 Instead of inducing new bone growth, some implants may cause resorption of bone because of the elements released into the surrounding tissue. Klawitter and Hulbert37 reported that even a slight hydration of calcium aluminate depressed its osteogenic effect. Thus, it may be that osteogenesis is depressed by using a soluble ceratiic material with an unstable surface.

Fabncabilrty and bioinertness

1700 1650 1850 1600 2996

Zr Ti-AI-V (c) Ti-Mo (d) Ti-Zr (e) Ta

“Bioinertness; (a) Venable CS (1937), Kawahara H (1973); Kawahara H (1980); (e) Kawahara H, Miura I (1984).

1800

1300-I 400 1250-I 350

IO-15 IO-15 IO-15 50-70 80-I 00 80-I 20 220-280 320370

(b) Linkow

H (1965);

*xx *** c* * XXX ***

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Bioinertness

LI (1968). Kawahara

Casting Castina Welding Sinterina ‘Welding Sintering Sintering Sintering Welding

Injection

Selfcuring Selfcuring Selfcuring Heatcuring Heatcuring Heatcuring Injection Injection

Fabrication

and metallic materials

Molding temperature (“C)

of polymeric

Ti (b)

Metals andalloys Co-Cr-Mo (a) Fe-Ni-Cr

Plastics Polymethylmethacrylate (composite) Tetramethylolmethane (composite) Silicone Polymethylmethacrylate (dental) Polyurethane Segmented polyurethane Polymethylmethacrylate Polysulfone (composite) Polysulfone Polyethylene (HDPE)

Materials

Pable 1

(c) Aragon

PJ, Hulbert

293340

30-230 (deform)

196294 (deform) 98-196 (deform)

2.45- 31.36 88.2 -102 114.36-l 35.24 89.2 -127.4 29.4

104.65

25-l 88 88-90

Bending strength (MPa)

SF (1972); (d) Miura

117

235 205

0.03 -0.17 0.3 -1.9 2.1 -5.6 4.96-4.80 0.49

2.3

2.45.8 6.0-6.5

Modulus of elasticity (GPa)

I,

Fabricability and bioinertness

(c)

1500 1500-2000 1800-2000 2000 1700-I 800 2050

900-l 200

1050-I 450

Fabrication temperature (“C)

of ceramics

Sintering Sintering Sintering Sintering Sintenng Sintering Melting

Melting Sintering Sintering

Fabrication method

***xx

xx**

*x*x

*a*

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x*x*

<‘*

29.4-186 0.4 3.2-12.7 6-l 6

68.6205 508 780 490 490 372 1274

105-215

**xx

137-150

***

98

Bending strength (MPa)

xxx

xxx

Bioinertness

10.8-17.6 0.09-0.19 41-56 12-18

16.8-27.4 50-70 225 294 392 372 392

41.2-121.0

98

Modulus of elasticity (GPa)

“Bioinertness; (a) Hench LL (1973); (b) Bfencke BA, BrBmer H (1975). Pernot F, Zarzycki J (1979). Yamamuro T et a/. (1980); (c) Mouroe EA et al. (1970). Aoki H et a/. (1976); (d) Dumas M et a/. (1974); (e) Bokros JC (1972); (f) Hulbert SF (1973). Nagai N et a/. (1982); (g) Hulbert SF (1970). Griss P (1978); (h) Hulbert SF (1970); (i) Sandhous S (1965); (j) Kawahara H (1975).

Hard tissues Cortical bone Cancellous bone Enamel Dentine

Bioinert ceramics Vitreouscarbon (d) LTI carbon(e) Zirconia (f) Silicon nitrate(g) Silicon carbide (h) Polycrystal alumina (i) Singlecrystal alumina(j)

Hydroxyapatite

Crystallized glass(b)

Bioactive ceramics Bioglass (a)

Materials

Table 2

Bioceramics for hard tissue repIacemmh

Figwe

3

187

Sapphire dental implants

In summary, bioactive ceramics are not suitable material for use in artificial dental root or the reconstruction of bone and joints, due to the low level of mechanical strength and the continued reduction of strength after long implantation times.

resorbable ceramic, consisting of P-Ca, (PO,),, beta-phase tricalcium phosphate ceramics (pTPC) was developed43+‘5; when implanted in bone, it is invaded with new bone ingrowth and is in turn resorbed.46,47 Favourable results were obtained with this material in repairing maxillofacial bone defects resulting from avulsive w~~~ds.~~ Usage of granular and powdered p-TPC for bone reconstruction and/or augmentation has been conGraves el’LzZ.~~ considered the use of soluble cal- sidered and meaningful results have been obtained from animal experiments and clinical evaluations. cium aluminate which dissolves in living tissue; the intention was to utilize the concept that im- Many studies using rats, dogs and monkeys”g-51 planted materials should act as a temporary scaf- have reconfirmed that P-TPC is a biocompatible fold to be replaced by the host bone. Bhaskar et matrix for bone regeneration and effective in bone reconstruction or augmentation in any form: mulal. ,39 Cutrigbt et aL40 and Getter et aL41have had success with the clinical impiantation of calcium ticrystalline, porous, powdered and/or granular. phosphate ceramics in bone and Driskell et aZ.42 the histological examinations have shown that the P-TPC implants are resorbed and concomitantly had favourable results with the use of tricalcium replaced by normal bone tissue. phosphate for the repair of cranial and alveolar Metsger, Driskell and Paulsruds2 reported clinbone defects. In 1974, the Research Project on Bioceramics, sponsored by the US Army Medical ical observations on the effectiveness of p-TPC Research and Development Command, had as to repair marginal periodontal bone defects. No adverse experiences were attributable to p-TPC. their major goals: (I) the development of ceramics for hard tissue repair; and (2) the fabrication and This clinical evaluation indicates that b-TPC is equivalent to autogenous bone tissue for repairsurgical implantation of porous-ceramic artificial ing marginal periodontal bone defects. tooth roots. In these two programmes, a porous,

188

H Kawahara

In recent years, two companies in the United States have produced new materials of l3-TPC called Augmen (Miter Inc.) for bone augmentation, and Peri-oss (Miter. Inc.) and Synthograft (Johnson & Johnson) for periodontal bone reconstruction. The granules are for packing into the bone defect in a periodontal lesion and the powdered particles are mixed with a physiological saline and biocompatible viscous polymer to form a paste; again for defect filling. Such types of biodegradable materials can be regarded as tissue organizers from the standpoint of an embryological technical term. It is, however, difficult to develop new biomaterials with tissue Bone formation around organizing factors. endosteal implants mainly depends upon the physical structure rather than any chemical factor. Resorbable ceramics of /3-TPC may act as a supporting structure for osteogenesis. In fact, there have been many reports which have shown new bone growth around many bioinert materials of ceramics, metals or plastics. Biodegradable materials which have slow kinetics of resorbability and nontissue irritability may be suitable and produce bone regeneration, and the resorbed area may be filled with new bone growth without any dead space. It is necessary to control the resorption rate so that it is equal to or less than ingrowth rate, so that the ceramic implant remains reasonably immobile to allow such ingrowth. Although soluble elements from P-TPC in bone may accelerate osteogenesis, this process is more likely to be influenced and depressed by the physiological status of the bone tissue. From these concepts, it is interesting to observe the studies of Plaster of Paris for the reconstruction of bone tissue which have been reported from 1892 to 1985. The first trial test was carried out in orthopaedic surgery by Dreesmann.53 Alderman,54 Shaffer and App,5s Be11,56 Calhoun, Green and Blackledges and Bahn58 used Plaster of Paris to repair alveolar bone defects. From the results in their trial, it was revealed that Plaster of Paris is a usable material for bone defect repair. A Yamaoka and Y Mikami have reconsidered the use of Plaster of Paris for bone augmentation and they have carried out histopathological examinations on the bone formation around the a-CaSOd 2Hz0 which showed a more reliable new bone formation than pCaSO, 2H,O. Our research group conducted trial

tests in clinical cases of alveolar bone defects and this reconfirmed that Plaster of Paris was a usable material for bone defect repair. However, those biodegradable materials still do not have sufficient mechanical strength. Also, excess absorption of bone tissue may be caused by the biodegraded elements instead of inducing new bone growth. It is not necessary for biodegradable ceramics to be suitable for all cases of filling bone defects, because it seems that osteogenesis around the biodegradable ceramics may be easily influenced by the physiological conditions of the bone tissue. Furthermore, it is unclear what substance promotes osteogenesis. It is reasonable to assume that the most reliable implant material would be extremely stable and insoluble in living tissues. Two Japanese companies have developed new nonbiodegradable ceramics of hydroxyapatite for bone defect repair, which are named Apaceram-G (Asahi Kagaku)59,60 and Bonetight (Mitsubishi).61 There are, however, problems associated with particle size and form. Fujisak?j2 had reported that nonbiodegradable alumina particles smaller than 0.6p.m were phagocytosed by mononuclear giant cells, polymorphonuclear leukocytes and fibroblasts. Alumina particles with a particle size range between 1-4pm were phagocytosed by mononuclear giant cells only. No phagocytosis was observed of the particles over 6pm which were covered with living cells. In vitro observations of the implanted hydroxyapatite granules confirmed that these findings were also applicable. Close cell adhesion to the surface of high density hydroxyapatite particles with large size particles of over 30E,cmwas seen, while small size particles of under 10pm were actively phagocytosed by giant cells spontaneously transformed from L-cells63 (Figure 4). Another problem with the new hydroxyapatites is that the particle form may have sharp edges and may produce mechanical irritation and act as a carcinogenic factor after long term implantation36,64,65(Figures 5 and 6). Tissue adhesion to bioinert ceramics

In the light of all these findings, the only materials suitable for the construction of artificial tooth roots and the replacement of bone are those which are chemically inert in living tissues and have

Bioceramics for hard tissue replacements

189

190 H Kawahara

Figure 5 Fibrosarcoma Induced by mechanical Irrit,ation with a square plate of sapphire: IS: implant space; X: fibrous capsule

Figure6

Adenofibroma

induced by mechanical irritation with a round

plateof polycrystal alumina

Bioceramics for hard tissue replacements

sufficient mechanical strength. The stable ceramics such as single crystal alumina, high density synthetic hydroxyapatite,@j silicon nitride and zirconia exhibit a high degree of biocompatibility and excellent adhesion to tissue and are useful materials from which to form artificial roots, bone and joint prostlleses.66 The surfaces of these ceramics absorb water, giving an aqueous layer which cannot be removed even in a vacuum at a temperature of 400°C (Figure 7). The slurfaces are not only stable, but also nonirritating, noncytotoxic, noncarcinogenic and superior in adhesion to the surrounding tissue.

0/0HiiO2

H’

SW-fact?

of

alumina

Crystal

structure

of alumina

Figure 7 A model for water absorption crystal alumina

on the surface of

When such ceramic beads were placed in a cell suspension of Hp cells and agitated, a monolayer of cells closely adhered to the surface after two days. After three days, the tissue layer built up on the surface of ceramic beads, with cell to bead and cell to cell adhesion. After four days, the ceramic beads were bonded together with strong cell to cell adhesion2 (Figure 8). If the glass beads were coated with oily materials, no cell aggregation was observed. Another interesting factor was investigated using culture techniques. When a plate of verre dure was placed in a MEM suspension of bone marrow cells derived from monkey femur, the cells showed cytoplasmic stretching due to their high wettability to the glass plate and adhered closely to the plate with a strong adhesion of 3.7 dynes/cm2 for 10 minutes.67 This was measured using viscometric methods.68 If the ceramic plate was contaminated with greasy materials, the cell adhesion was disturbed and a decrease of adhesive strength resulted, therefore, it is important that all implant surfaces should be kept clean during the operation (Figure 9).

The mechanism

191

of tissue adhesi

From observations in vitro and in vivo, it appears that the following three steps represent the interfacial reaction between the stable ceramic implant and surrounding tissue6 (Figures 10 and ll), Step I: Immediately after the insertion of the ceramic implant into bone tissue, the implant surface is coated with a blood clot. Both epithelial and connective tissues are bound to the aqueous layer on the implant surface through the sandwich layer,3*4 consisting of blood clot, proteins, lipids, polysaccharides and proteoglycans between the ceramics and cells. This layer is probably involved in establishing the strong adhesion between the tissue and the implant which develops over time. Step 2: One month after insertion of the implant: the blood clot layer is reorganized as the phagocytes and fibroblasts reach the implant surface and come into direct contact with the surface via cytoplasmic extensions. Gingival epithelial cells, if they retain their vitality, may absorb the water in the sandwich layer and syneresis of that layer may create a good adhesion between the implant surface and the epithelial cells. Step 3: Three months later, the collagen fibres bind directly to the implant lying vertically to its surface. This is unlike the fibres resulting from a foreign body reaction which lie parallel to the surface. Gingival epithelial cells adhere to the implant surface and may produce hemidesmosomes in their unit membrane. Alumina

and rirconia

Polycrystalline alumina and single crystal alumina, which have high mechanical strengths compared with those of the other stable ceramics, allow delicate, complex designs to be fabricated. The first clinical application of single crystal alumina was reported by Kawahara and an American patent and Japanese patent exist.24,25The biocompatibility of crystalline alumina (sapphire) has been proven by tissue culture and animal experiments by Kawahara and coworkers since 1972. Kawahara and his coworkers developed various types of sapphire implants which have been used to fabricate artificial dental roots2.s (Figure 3), bone screws16 (Figure 12), bone and joint prostheses (Figure 14) with superior wear resistance,

Figure 8 Cell adhesion to glass beads by gyratory cultivation: days cultivation

Figure9

Cell adhesion

to a single crystal plate (left),

(I) two days cultivation,

(2) three days cultlvatlon;

(3) four

no cell adhesion to a elate coated with sillcone

irlghti

excellent biocompatibility and close adhesion of fibroblasts, collagen fibres and bone tissue6” (Figures 15 and 16). The adhesive strength at the interface of sapphire with bone was measured with a torquemeter one month, three months and six months after the insertion of a sapphire screw into the mandible of a monkey. The adhesive strength doubled after three months. The adhesion was verified by

electron microscope observations which showed partial ankylosis and attachment of new bone to the implanted screw.” Zirconium oxide ceramics have excellent physical properties for dental implants and bone and joint replacements. Nagai et al. reported that histopathological observations by LM and EM demonstrated close adhesion between the endosteal implant of zirconia and the adjacent new

Bioceramics

for hard tissue replacements

STEP-

STEP--BLOOD CLOT LAYER

Figure IO

Interfacial reaction between

COLLAGEN FORMATION

bioinert ceramics and surrounding

bone22,23 (Figures 17 and 18). A direct bond was seen without any soft tissue at the ceramic/tissue interface. inflammation

and carcinogenicity

Ceramics generally have a high Young’s modulus compared with that of bone and tooth hard tissue. The difference between the Young’s modulus of the implant material and the surrounding tissue produces biomechanical problems and movement at the interface. Ceramics with higher Young’s moduli generally induce more irritation responses in the adjacent tissue. 66 Efforts are being made to develop a new implant material which is similar in its mechanical properties to the tissue being restored. At present, there are no ceramics which have adequate biomechanical compatibility, well balanced in terms of strength (tensile, compressive and bend), Young’s modulus and elongation to

I93

NEW BONE - CERAMICS INTERFACE

bone tissue

Failure. Few cases of bone resorption induced by biomechanical inbalance have, however, been experienced in the T-types and U-types of dental implants3 (Figure 19) made of single crystal sapphire with a high Young’s modulus of 392 GPa. In experiments considering the carcinogenicity of chronic mechanical irritation, various forms of sapphire, round plate (Smm), square plate (8 x 8mm), round plate (8mm) with perforated pore (lmm) and porous plate (pore size 50-50Opm, porosity 28-32%)) were implanted into the subcutaneous tissue of 600 female four-week-old Sprague-Dawley rats (CLEA). After two years implantation, the carcinogenicity was investigated by histopathological methods (Figures 5 and 6). Fibroma (10.5%) and fibrosarcoma (15.8%) in the square plate; fibroma (5.2%) and fibrosarcoma (15.7%) in the round plate; fibrosarcoma (16.6%) in porous plate and no tumour formation in the perforated round plate were observed (Table 3). The carcinogenicity of these neoplasmic

194

H Kawahara

Figure II(a) Change In the hrstologlcal reactron atthe implant-bone Interface with a lapseof time after the Implantation of sapphire in dog mandible, 63twoweeks Implant Blood clot layer (bl) over the fibroblastic network (f): IS: implant space.

Figure II(b) Implant at three weeks, partof blood clot layer disappears ( 7 1.Frbrobiastrc net work 1”) between the implant and bone (bo) is established with fibres of vertical drrectron to the implant surface

( T ) between the implant and

ure II(c) Implantation new bone tissue (bo)

after three months, close adhesion

Figure II(d) bone damage

after 12 months under the functional condition with biting load,

lmplantatlon

( T t ) and resorption ( t ) were observed at the interface and lntrabony area

Figure 15 18 months the implant. denoted by

The free grngival margrn from a sapphire implant after demonstrates the crevicular epithelrum interface with The apical termination of the crevicular epithelium is the arrow (by courtesy of Dr McKinneyP

Figure 16 A sapphire implant In functron for 12 months. The bone interfaces directly with the implant and has regenerated superiorly over the shoulder (by courtesy of Dr McKinneyP

Bioceramics for hard tissue replacements

Figure 12

197

Bone screws and bone tapper made of single crystal alumina for the fixation of fractured bone’6

tissues was reconfirmed by intraperitoneal injection. From these findings, it is presumed that the carcinogenic response is induced by chronic mechanical irritation with the sharp edge of an implant having a high Young’s modulus. These findings demonstrate the importance of designing an implant without any sharp edges. It must be emphasized, however, that there have been a number of successful implants of bioinert ceramics reported in clinical cases with no carcinogenic tissue. This may be due to the circumstantial difference between subcutaneous soft tissue and the hard tissue of bone and also species differences between rats and humans.

It is important to ensure initial retention and fixation of the implant in bone tissue in order to prevent mechanical inflammation and the formation of collagen fibre encapsulation. Movement

of the implant induces inflammation and disturbs the normal formation of collagen and new bone growth and thus causes loosening of the implant. In order to keep rigid retention it is important to consider the relationship between the relative volume of implant and surrounding bone tissue. The more bone tissue retained after the implantation, the better are the chances of ensuring a reliable fixation. Consequently, small implants (which allow more bone retention) can and should be constructed of rigid, high strength materials. Many cases of fracture in conventional ceramic screw dental implants were experienced in animal experiments and with the clinical use of both selftapping or screws into the bone, as well as under the impact stress of body movement and mastication. Sapphire has produced favourable results in intricate dental implants in spite of its high Young’s modulus. The small size of sapphire implants tend to induce stress concentration in the bone tissue surrounding the implant. However, this stress concentration may be resolved

198 H Kawahara

Figure 13

Table 3

Sets of artificial bones and joints made of single crystal and polycrystal aiuminat6

Neoplastigenicity

induced by mechanical

irritation of AlzO,

Neoplasm

Fibroma (%)

Fibrosarcoma (%)

Test piece

8mm porous

0

Poly

16.6

8mm

Single Poly

0

0

0

0

8mm

Single Poly

10.0 5.2

0 15.7

8x8mm

Single Poly

10.5 10.5

5.3 15.8

Bioceramics for hard tissue replacements

199

Bioceram Single Crystal lumina

8 ioceram Polycrystalline Alumina Stainless ASTM CoCr-Mo

Steel 316

Alloy

Titanium Cortical

IC

0mpressivE

w

Flexural

!Z?Z4

Tensile

Bone

(Femur);

I

5.000

t

10.000

I,

15. 000’50. ccc kz

Figure 14 metals)

effectively

cm’

Comparison of strength of various biomaterials (“0.2% yield strength is shown for the

by using two

piece

screw implants6

(Figure 201, instead of large monolithic implants with low mechanical strength such as hydroxyapatite, vitreous carbon and polycrystalline alumina. To avoid the fracture of dental screw implants the following sizes should be used: 5 mm or more in hydroxyapatite and vitreous carbon, 3.2mm or more in polycrystalhne alumina and 2.2mm in sapphire screw implants. It can be seen that it is possible to retain more bone in the case of sapphire implants, which leads to an excellent prognosis for this implant.

Porous prosthetic implants of metals, plastics and ceramics create reliable fixation of the implant by tissue ingrowth into the pores. The first studies on the speed of tissue ingrowth into various kinds of porous ceramics were carried out by Klawitter and Hulbert39 and Hulbert et a1.70 Selting and Bhaskar’r measured the structural strength of the interface between bone and nondegradable porous ceramics of Al,@ and concluded that normally

healed cortical defects 10 weeks postoperatively are about 80% of the uninjured cortical strength. Predecki, Stephan and Auslaender, Mooney and Kirkland72 investigated the dependency of pore channel diameter on bone ingrowth kinetics using alumina and titanium. They reported that the most rapid bone ingrowth was obtained with 500, 52.5 and 1000pm diameter channels in all the titanium and alumina samples. However, the fraction of any given channel crosssection filled with bone was observed to decrease as the channel diameter increased. The presence of interconnecting fine porosity was found not to be essential for bone ingrowth, nor does it have a major effect on ingrowth kinetics. Clinical radiographic, microradiographic and histological examinations of porous alumina implants into the mandibles of monkeys revealed ingrowths of fibrous tissue and bone tissue into the interconnecting channels. No adverse tissue reaction was found in porous alumina by Pedersen, Haanaes and Lyng.73 Studies in vitro using L, HeLa and bone marrow cells, demonstrated their rapid ingrowth into the porous alumina.74 No cell ingrowth was observed at distances more than 1.2mm deep in materials

200

H Kawahara

Figure 17 Zirconia implant contacts closely to the surrounding bone tissue in a monkey’s mandible after four months implantation (by courtesy of Dr Nagai)23

Figure 18 No soft tissue capsule at the interface of zirconia implant and bone tissue, including a bone cell (by courtesy of Dr Nagai)z3

with small pore sizes of 20,um or 50pm, but in materials with large pore sizes of 135-150pm, nine and 32 ingrown cells were observed deep into the surface. Cells could not retain their normal cell activities when the pore size was less than 20pm. Again, this may depend upon the high wettability of cells on the surface. With larger pore sizes, ranging up to 500pm, large numbers of ingrowing cells were observed. The speed of cell ingrowth depends upon the cytoplasmic locomotion and cell mitosis and the former predominates in the depth of porous materials.

alumina implants reinforced with a single crystal core were designed to prevent fracture under biting stresses. 75 New bone growth into porous alumina has produced successful fixation of implants replacing a single standing tooth after one or two months implantation in the mandibles of dogs and monkeys. Light and electron microscope studies have demonstrated the interlocking ankylosis between the implant and bony ingrowth. Close adhesion of the epithelial tissue and bone to the post of single crystal alumina was also observed. l6 These findings obtained in cell culture animal experiments and clinical cases suggest porous alumina implants reinforced with a single crystal core can be highly successful as a dental and orthopaedic implant. However, appliances of porous alumina, which

Porous alumina implants

The mechanical strength of porous alumina with pore sizes of 50-500pm is not adequate. Porous

Bioceramics for hard tissue replacements

Figlure19 Bone resorption impiants of sapphire

around the T-type and U-type

have a high Young’s modulus, should be designed so as to avoid possible neoplasmic responses following prolonged mechanical irritation of the tissues by sharp edges. Because of this biomechanical problem, Homsyr6 developed Proplast which has a low Young’s modulus. It is a porous material consisting of Teflon (polytetrafluoroethylene) and glassy carbon fibres. Many successful implants have been made by coating the surfaces of metallic and ceramic implants with Proplast. Proplast does, however, have some drawbacks, because polytetrafluoroethylene has a low wettability to cells and does not produce close adhesion between tissues and its surface. A greater success rate for porous implants can also be achieved by using a buffer layer of segmented polyurethane in the superstructure in order to break the impact stress during mastication3s4 Five kinds of segmented polyurethane

201

ranging from soft to hard have been developed and named ‘Biotron’. They have a simil,ar viscoelastic behaviour to natural periodontal membrane (Figure 22). Suitable Biotron can be selected according to requirements from the five different Young’s moduli of Biotron. The continued success with implants may depend upon achieving rigid retention by using the porous forms of bioinert alloys or ceramics with high Young’s moduli which are free from sharp edges. A buffer layer or stress absorber shou!d be added to the superstructure in order to break the concentration of stress within the alveolar bone. Kawahara,” Hirabayashi and Kawahara,77 and Yamagami and Kawahara78 have developed a porous alumina material reinforced with a single crystal core of sufficient strength to be used for artificial dental roots (Figure 21), bone replacement and joint prostheses. Rigid fixation can be obtained by macroscopic interlocking between the porous implant and new bone. Porous alumina implants (which have a high Yo.ung’s modulus) should be used conservatively, since a failure of the bone at the bone-implant interface due to biting stress during mastication is possible. To protect the implant from the biting stress, the method of fibroosseous integration was considered by Weiss.7” However, periimplantium may introduce loosening of the implant. It is important to control geometric distribution of bone tissue, osteoid tissue and fibrous tissue at the bone-implant interface because the implant must be supported with a biologically dynamic interface in order to produce a successful porous implant. Porous implants with multipore size at the bone-implant interface have resulted in many successful implants because of these qualities. References

Kawahara H, Ochi S, Yamagami A. Biological testing of dental and surgical materials by means of tissue culture. Second Proceedings of the International Academy Pathology 1965: 79.

of Oral

Kawahara H. Future vision of implantology. In: Kawahara H ed. Implantology and biomaterials in stomatology. Ishiyaku Publ, 1980: 1-17.

202 H Kawahara

Figure 20

The bone fracture of a condyle head fixed with two sapphire bone screw.9

Kawahara H. Cellular response to implant materials. Inter Dent J 1983; 33: 350-75. Kawahara H. Materials for hard tissue replacement. In: Lin OCC, Chao EYS eds. Perspective on biomaterials. Elsevier Sci Publ, 1986: 167-206. Kawahara H, Yamagami A, Nakamura M. Biological testing of dental materials by means of tissue culture. Znter Dent J 1968; 18:443. Kawahara H. Materials for hard tissue

replacement. Inter J Oral Implant 1985; 3: 17-27. 7 Kawahara H. Cellular responses to dental materials by means of tissue culture. J Jap Sot Dent Apparat Mater 1962; 3: 105. 8 Kawahara H. Biological evaluation of implant materials - cell adhesion to material. Transactions of the 10th Annual International Biomaterials Symposium, San Antonio 1978; 2: 11-15.

Bioceramics for hard tissue replacements

Figuse 21

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Five types of porous alumina implants relnforced w&h sapphire core posts

4

Figure 22 Crosssection of a porous alumina implant with a sapphire core post and a stress absorber with a Biotron layer in the superstructure

9 Kawahara H, Nakamura M, Yamagami A et al. Electron microscopic studies on the cellmetal interface in vitro and in vivo. Transactions of the 11th Annual International Biomaterials Symposium, Clemson, 1979. 10 Kawahara H, Nakamura M, Takada S. Toxicity and solubility of nickel from Ni-Cr alloys in tissue culture. Transactions ofthe 7th International Biomaterials Symposium, San Diego, 1985: 157. 11 Kawahara H, Nakamura M, Ishizaki N, Takeda S. Cytoxicity and solubility of cobalt from Co-Cr binary alloys in tissue culture. Transactions of the 8th International Biomaterials Symposium, Orlando 1986; 12 Inoue T, Cox JE, Pilliar RM, Melcher AH. Effect of the surface geometry of smooth and porous-coated titanium alloy on the orientation of fibroblasts in vitro. J Biomed Mater Res 1987; 21: 107-26. 13 Kawahara H, Imai K, Miura I, Okuno 0. Cellular responses to new bio-alloys of titanium for porous structure, in vitro. Transactions qf the 15th International

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