Biomaterials 23 (2002) 2569–2575

In vivo behaviour of hydroxyapatite coatings on titanium implants:

a quantitative study in the rabbit

G.L. Darimont

a,

*, R. Cloots

b

, E. Heinen

c

, L. Seidel

d

, R. Legrand

a

a

Department of Oral Implantology and Periodontology, Institute of Dentistry, University of Liege, CHU, Sart-Tilman 4000 Liege, Belgium

b

Laboratory of Inorganic and Structural Chemistry, University of Liege, Domaine Universitaire Sart-Tilman 4000 Liege, Belgium

c

Institute of Human Histology, University of Liege, rue des Pitteurs 4020 Liege, Belgium

d

Institute of Biostatistics, University of Liege, CHU Sart-Tilman, 4000 Liege, Belgium

Received 25 May 2001; accepted 20 November 2001

Abstract

The aim of this study was to evaluate quantitatively the behaviour of in vivo hydroxyapatite coated implants (HA) in the rabbit

over time, and to compare the results with observations made on titanium plasma spray implants (TPS). Results were analysed
according to the percentage of bone contact. Eighteen HA cylindrical implants (3.25 8 mm) and 6 TPS cylindrical implants from
Steri-Oss were placed in the epiphysis of the femur in 24 white rabbits. Each rabbit received one implant. Three rabbits with one HA
implant (n ¼ 3) and 1 rabbit with one TPS implant (n ¼ 1) were sacrificed after implantation periods of 2, 4, 6, 8, 10 and 12 months.
Implants were cut along the long axis and prepared for histological and histomorphometrical evaluations. Measurements of coating
thickness and percentage of bone contact were performed with scanning electron microscopy analysis on the sides of the implant, in
3 different types of bone, namely cortical, trabecular and marrow. In cortical bone, dense bone was apposed to the HA implants:
from 92.3

75.5% at 2 months to 89.676.5% at 1 year, with no significant regression of HA thickness (P ¼ 0:37). TPS coating

showed less bone contact, but thickness was stable (P ¼ 0:46). In trabecular zone, where bone contact was less pronounced, a
significant regression of HA coatings thickness (P

o0:05) was observed. Nevertheless TPS coatings were stable (P ¼ 0:81).

Histomorphometrical results demonstrated that a highly significant regression (P

o0:0001) of HA thickness was observed in the

marrow area, where the bone-to-implant contact never exceeded 7.6% from 2 to 12 months. TPS coating did not reveal any sign of
resorption (P ¼ 0:88), despite a rare bone contact. Histological analysis revealed inflammatory and giant cells, principally in the
marrow area in contact with HA coating, but always in restrictive numbers. We conclude that bone contact protected the HA
coating from resorption. r 2002 Elsevier Science Ltd. All rights reserved.

Keywords: Bone-implant contact; Hydroxyapatite coating; Titanium; Endosseus implants

1. Introduction

Implantation Techniques have been firmly established

as an integral part of oral rehabilitation treatment. A
number of implants with different surfaces have been
used over the last 30 years, with the primary objective of
increasing both the quality and quantity of bone contact
[1–3]. This has provided the rationale for using
hydroxyapatite (HA) coated implants [4–6]. Their
roughness increases surface contact 5 times as compared
to smooth implants, and their chemical composition is
similar to the mineral fraction of bone. They have been
widely used and valued in oral implantology and

orthopaedicsurgery for over 20 years mainly in Europe
and to a lesser extent in USA. We observed a faster and
better bone contact than for the titanium implant (Ti)
[7–10]. Strong criticism has questioned the long term.
Degradation of the coating appears progressive, leaving
the Ti substrate without any bone contact [11].
Similarly, many histological studies have indicated that
the initial gain in bone contact diminished over time,
ending up with a lower value compared to Ti implants
[12,13]. These factors lead to implant mobility that
could lead to a break down in bone contact in peripheral
areas.

Resorption–dissolution of the coating is often seen in

failed cases, but is this the cause or the consequence of
the observed mobility? We thought it would be
appropriate, where it is, to quantitatively evaluate the

*Corresponding author. Tel.:/fax: [phone].

E-mail address:

[e-mail] (G.L. Darimont).

0142-9612/02/$ - see front matter r 2002 Elsevier Science Ltd. All rights reserved.
PII: S 0 1 4 2 - 9 6 1 2 ( 0 1 ) 0 0 3 9 2 - 1

behaviour of HA coating shielded from any external
influence either chemical or physical. In this respect, we
have quantitatively analysed the evolution of the
thickness of commercially available HA coated im-
plants. By comparison, we analysed the behaviour of
TPS coated implants.

2. Material and methods

One Ti6Al4V cylindrical implant with a diameter of

3.25 mm and a length of 8 mm (Steri Oss Yorba Linda,
CA, USA) coated with a synthetic HA coating
(Ceramed Lakewood, CO, USA) was placed in the
distal femur epiphysis of the right thigh in 18 male and
female ‘‘White of Termonde’’ rabbits, aged 6 months to
1 year, with body weighing between 4 and 5 kg. Six other
rabbits received one titanium plasma sprayed implant
(TPS) (Steri-Oss) in the same anatomical location. The
rabbits were anaesthetised using 4 ml of xylazine at
2%(Rompun

s

), 10 ml of ketamine (Imalgene

s

), and a

local injection of 1.8 ml of articaine (Alphaca

.ıne Sp

s

).

The right thigh was shaved and washed with Isobeta-
dine

s

. Shallow and deep incisions were made with a

n115 blade. Drilling of the cortical bone was performed
in sterile distilled water, followed by the perpendicular
insertion of implant to the point of contact with
opposite cortical. Shallow and deep incisions were
closed with resorbable sutures (Vicryl

s

4.0, Johnson&

Johnson.). Wounds were cleaned with Isobetadine

s

.

The rabbits received post-op a one shot of 100 mg of
Novalgine

s

; there was no restriction of either move-

ment or food.

2.1. Experimental design and sample preparation

Three rabbits each with one HA implant and 1 rabbit

with one TPS implant were sacrificed by lethal injection
of phenobarbital at 2, 4, 6, 8, 10 and 12 months. A 3 cm
section of the femur of each rabbit, containing the
implant with the surrounding bone was removed, placed
immediately in 10% formaldehyde solution for 48 h,
dehydrated over 3 days in graded ethanol solution
(70%, 80% and 100%), brightened in toluene and
embedded in methyl-metacrylate resin purified in caustic
soda and then dehydrated in copper sulphate. Polymer-
isation was achieved with bazoilic peroxide. This
preparation gives the advantage of increased hardness,
and guarantees high sample quality with minimum
tearing, compared to commercial solutions. The samples
were mounted in epoxy resin. The block was cut along
the long axis of the implant in an Accutom (Struers,
Denmark) sectioning machine with a circular diamond
blade HV 500 at a speed of 150 mm/s. The specimens
were polished successively with Si C papers of decreased
roughness (220, 500, 1200 and 4000) during 45 s, under

water, at the speed of 150 revs per minute, with a load
pressure of 30 N. The final polish were obtained first by
polishing on a soft napped cloth with a 3 mm diamond
paste during 6 min under a load pressure of 30 N, and
then with a 1 mm diamond paste with the same soft
napped cloth during 1 min under the same pressure.

A preliminary evaluation under an optic microscope

enabled the choice of the sample best suited to scanning
electron microscope (SEM) analysis, out of respect for
the quality of the exposed surface area (free of tearing,
absence of bubbles in resin, etcy).

2.2. Histomorphometric quantification of bone contact

The 24 anatomicsamples were gold sputter c

oated

thickness of (20–40 nm) and analysed at different
magnifications in a Hitachi S 2500 SEM. The lowest
magnification (25 ) allowed correction of the section
with relation to the median axis. It was not possible to
perfectly align the centre of the implant; and so this
deviation was corrected using the Pythagorus theorem.
The analyses were done at a magnification of 800 .
A field of view of 136 mm was obtained at this
magnification. Four different observation sites were
chosen, two on the left and two on the right side of the
implant for each of the 3 different bone types. First, the
percentage of bone contact was measured all along the
window, on both sides of the implant, adjacent to the 3
different types of bone. Secondly, coating thickness was
measured in the same location, and with mean values
obtained from the minimum and the maximum thick-
ness. For the 3 HA implants (n ¼ 3) for the different
implantation periods (2, 4, 6, 8, 10 and 12 months), 2
different observation sites were chosen on both sides of
the implant for the 3 different types of bone. The same
procedure was performed for the TPS implants (n ¼ 1).

2.3. Histological and histomorphometrical evaluations

The analysis consisted of:

*

Light microscopic evaluation. A subjective descrip-
tion at different magnification was reported after
staining with toluidine blue.

*

Measurement of the amount of bone contact
percentage in the 3 different types of bone.

*

Observation of coating thickness of the HA and TPS
coatings over time in the 3 bone regions.

2.4. Statistical analysis

Results were expressed as mean

7standard deviation

(SD). Linear regression was used to analyse data over
time. Statistical calculations were done by using the SAS
6.12 (SAS Institute Inc., Cary, North Carolina, USA).

G.L. Darimont et al. / Biomaterials 23 (2002) 2569–2575

2570

All results were considered to be significant at the 5%
critical level (P

o0:05).

3. Results

3.1. Light microscopic evaluation

The animals appeared to be in good health through-

out the experimental period. No sign of inflammation or
adverse tissue reaction was observed around implants.
All implants were in contact with the 3 different bone
areas (cortical, trabecular and marrow). Light micro-
scopic evaluation showed a good integration of all
implants.

3.2. SEM observations

To compare the density of the HA coating of the SEM

observation of our samples, we have first realised a
polished cross section of a virgin HA implant (Fig. 1).

3.2.1. Bone contact percentage on HA and TPS coatings

Mean values and SDs of the amount of bone contact

on HA and TPS coatings, after implantation, are
displayed in Table 1. As indicated by the results, the
cortical site showed a higher percentage of bone contact
compared to trabecular bone and marrow area.

3.2.1.1. Cortical bone. HA coatings were almost com-
pletely covered with dense bone (Fig. 2). At 2 months,
we observed new lamellar bone apposition without any
infiltration of fibrous tissue. No inflammatory cells were
visible adjacent to the implant. Haversian remodelling
structures could be observed. Percentage of bone
contact reached maximum values after 2 months
(92.3

75%) and was not subject to any significant

modification (P ¼ 0:92) during the rest of the experi-

mental period. For TPS coating, progression of bone
apposition was different. Percentage of bone contact
increased significantly (P ¼ 0:0036) incrementally, from
57% at 2 months to 79% after 1 year.

3.2.1.2. Trabecular bone. With HA coating, newly
formed lamellar bone was observed in the healing
process, but in contrast with the cortical area, the

Fig. 1. Cross section of a polished virgin HA implant. Coating was
very dense, with several cracking lines (arrows). SEM 800 .

Fig. 2. Two months HA implant in cortical zone. Note the extended
bone contact (arrow) and the newly formed Haversian systems (h).
Lengthways section SEM 400 .

Table 1
Amount (%) of bone contact with HA or TPS implants, according to
time, in the 3 different bone areas

HA

TPS

Mean

7SD, n ¼ 3

Mean, n ¼ 1

Cortical bone
2 months

92.3

75.5

57

4 months

86.0

70.4

65

6 months

88.7

75.5

62

8 months

86.3

78.9

70

10 months

92.3

74

72

12 months

89.6

76.5

79

Trabecular bone
2 months

53.0

79.5

41

4 months

52.6

76.6

43

6 months

56.3

78.3

51

8 months

59.6

79.7

56

10 months

59.3

77

59

12 months

60.3

710

54

Marrow
2 months

5.3

73.5

8

4 months

6.0

72

7

6 months

6.0

73.6

8

8 months

7.1

73.6

10

10 months

6.6

73.5

9

12 months

7.6

74

8

G.L. Darimont et al. / Biomaterials 23 (2002) 2569–2575

2571

amount of bone apposition was less pronounced. For
the TPS coating, bone contact increased during the
experimental period and was statistically significant
(P ¼ 0:0234). Osteoclasts were visible occasionally in
the bone–implant interface.

3.2.1.3. Marrow area. Measurement of bone contact on
the HA implant was very difficult to perform due to the
separation of the coating into small aggregates. Bone
apposition was only observed in contact with the border
of trabecular bone, and never exceeded a percentage of
7%. Modification over time was not significant
(P > 0:05). The same could be described for the TPS
coating. Poor bone apposition, between 8% and 10%,
was observed. Polynuclear cells were observed princi-
pally around HA coating.

3.2.2. HA and TPS coating thickness

Mean values and SDs of coating thickness, according

to time after implantation, bone localisation and type
of coating (HA vs. TPS) are presented in Table 2.
A control test sample which had not been implanted was
analysed along the length and was found to have an
average HA thickness of 51.5 mm

71.92. In the cortical

area the SEM images showed a close contact between
HA and bone over time. The maximum thickness of HA
coating was subject to a very slight linear decrease,
which was not significant (P > 0:05) and with mean
values of 51.3 mm

71.69 at 2 months and 48.1 mm72.36

at 12 months; TPS values were stable. In trabecular
contact (Fig. 3), HA coating thickness was subject to a
significant linear decrease (P

o0:05), from 51.9

70.9 mm

at 2 months to 44.2

70.6 mm at one year. Thickness of

the TPS coating remained stable over time (Table 2 and
Fig. 4). In the marrow area (Fig. 5), non-linear evalua-
tion of the mean values of HA thickness appears
impressive (P

o0:05) and proved that HA thickness

decreases over time (Fig. 6). At 2 months, the mean
value is 49.2

71.5 mm. Over several months, the HA

coating was completely removed but remained adjacent
to the titanium surface, with a mean thickness of
18.2

75.7 mm. Standard deviations were larger, due to

the great disparity of the results.

Again, for TPS coating, the values remained perfectly

stable.

Table 2
Thickness (mm) of HA and TPS coating according to time in the 3
different bone areas

HA

TPS

Mean

7SD, n ¼ 3

Mean, n ¼ 1

Cortical bone
2 months

51.3

71.6

51.9

4 months

51.4

71.3

55.4

6 months

50.4

72.2

56.2

8 months

49.8

71.2

57.3

10 months

49.5

72.3

50.2

12 months

48.1

72.3

51.2

Trabecular bone
2 months

51.9

70.9

49.9

4 months

52.4

70.7

53.2

6 months

47.6

71.3

52.1

8 months

46.1

71.3

55.9

10 months

43.2

71.5

52.6

12 months

44.2

70.6

50.6

Marrow
2 months

49.2

71.5

50.5

4 months

45.7

70.9

47.1

6 months

43.5

71.1

49.9

8 months

41.2

70.7

46.3

10 months

36.0

73.0

48.1

12 months

18.2

75.7

48.4

Fig. 3. HA in trabecular contact after 10 months. Note a slight
dislocation of the HA coating (arrows). Lengthways section SEM
700 .

Fig. 4. TPS implant in trabecular bone after an implantation period of
8 months. The lack of bone contact (arrow) did not affect behaviour of
TPS coating. Lengthways section. SEM 360 .

G.L. Darimont et al. / Biomaterials 23 (2002) 2569–2575

2572

4. Discussion

The Aim of the present study was to investigate the

physical changes of an HA coating on a titanium
substrate. We compared it with TPS coated implants
placed in rabbit femur epiphysis. Evaluation of histo-
morphometrical data demonstrated a good HA coating
stability, principally in cortical bone, where a close bone
contact was observed. In contrast, a significant decrease
of HA thickness was observed in the marrow area,
where the lowest bone-to-implant contact was reported.
These findings confirm earlier observations. Authors
described a poor bone contact between the HA coating
of a total hip prothese and marrow of the human femur
[14]. Others observed a higher resorption rate of the HA
thickness in bone marrow in contrast with cortical or
trabecular bone [15,16].

In the case of TPS coatings, neither the amount of

bone contact nor the implantation period modified the

thickness of the coating. The different measurements
showed that a more complete bone contact preserves the
HA coating, thus this indicates that bone protected the
HA coating from degradation adjacent to cortical bone.
Also, light microscopic evaluations revealed that bone
formation proceeded faster and resulted in more bone
apposition for HA implants than for TPS implants. In
general calcium-phosphate coatings, whatever their
composition, promote bone apposition and differentia-
tion of mesenchymal cells to osteoblasts, however, the
exact role of these ceramics on bone healing is still
unknown. Several suggestions have been proposed to
explain the mechanism:

*

increased adsorption and production of proteins
(growth factors);

*

easier adhesion of osteoblasts on a recognised surface
(HA is the mineral part of the bone);

*

promoted influence on centrifuge osteogenesis (from
implant to bone) and

*

increased crystal growth and matrix mineralisation.

HA coatings not only provided an earlier stabilisa-

tion, but also close contact with bone, without
interposition of fibrous tissue. Some fracture lines
appeared at the coating–implant interface, and between
the tissue and the coating where no or little bone was
present (Fig. 7). The fracture line was occasionally
observed in the bone, indicating that bonding between
bone and HA coating is stronger than between implant
and HA coating. In contrast, fracture was never
observed between titanium and TPS coating. Surface
roughness also influences bone apposition [7,17]. For a
non-coated titanium implant, bone apposition never
exceeded 60% [18]. Although the HA coating displays a
small decrease in the long-term bone bonding, this
experiment revealed more than 90% contact of cortical
bone to the HA coating and 79% for TPS. Resorption in

Fig. 5. HA implant in marrow bone after 10 months. Dissociation of
HA coating in small particles (white arrows) becomes more significant
over time. Lengthways section. SEM 850 .

Fig. 6. Minimum values over time of the HA and TPS coatings in
contact with bone marrow Regression is highly significant (P

o0:001).

Fig. 7. Where we observed a fit bone contact like here in the trabecular
bone, the cracking line always appeared between HA coating and Ti
(double arrows), never between bone and HA (simple arrow).
Lengthways section. SEM 130 .

G.L. Darimont et al. / Biomaterials 23 (2002) 2569–2575

2573

areas of poor bone quality contact was evident. The
question remains as to the behaviour after more than 1
year. One study indicated that the resorption of the HA
and related calcium-phosphate coatings occurs in the
amorphous calcium-phosphate regions of the coating
and stops where the coating has 100% of crystallinity
[19]. Thus high crystallinity protected HA against
resorption. It is interesting to note that in our
experimentation, implants with percentage of coating
crystallinity between 63% and 94% were used. Further-
more, plasma spray technique could modify the original
titanium surface texture [20]. Another way to solve this
problem would be to promote the use of another deposit
technique to obtaine a thin layer of HA coating of the
order of 1–5 mm, instead of 50 mm obtained with the
classical plasma spray technique [21]. In fact, several
studies have already demonstrated that coating thick-
ness has an influence on bone response [22]. A thin layer
of HA obtained with other techniques than plasma
spray, such as the ion beam sputter technique [22]
seemed to induce good bone apposition, no alteration of
original titanium surface, and a sufficient resorption
time of the HA to allow a progressive bone contact on
titanium substrate.

But the resorption we observed appeared to be

dictated by a physiological process, and not a patholo-
gical process. The disparity of the values taken from the
same sample over time could explain the influential role
of the coating quality, and of the direct surrounding
area with respect to the resorption–dissolution rate. In
the marrow, the poor bone contact and the presence of
abundant fluids and cells are probably a sign of high
enzymaticac

tivity. It is thought that the degradation

process of the HA coating could be mediated by
monocytes [23,24], and to a lesser degree by polynu-
cleated cells, stimulated by a series of mediators. These
latter mediators may be cytokines HILDA/LIF [25],
ginterferon, hormones such as vitamin D3, Insulin-like
growth factor I [26], transforming growth factors [27]
proteins such as osteocalcine and osteopontine [28],
vitro and fibronectine, acids such as l-glutamicac

id,

which acts as a stimulator for monocytes and poly-
nuclears, or the antibiotic polymixin B which act as
inhibitors [29]. Others have proven that the mineralisa-
tion on the coating surface is influenced by the protein
adsorption from the surrounding environment [30], but
that the composition of the latter is itself dependent on
the nature of the coating. It seems evident that the
ceramic quality influences nearby cell activity, changing
primary cells into osteoblasts or osteoclasts [31–33].
Other authors believe that the dissolution of the HA is a
necessary preliminary stage in the transformation of
biological equivalents that have a mediator role between
osteoclast and osteoblast differentiation [34].

However, HA coatings differ from one manufacturer

to another, and even those coming from the same

vendor show differences [35]. Percentage of crystalline
vs. amorphous phases of the powder, Ca/P and HA/
TCP ratio, percentage of macro and micro porosity [36],
thickness of each layer, final thickness of the coating,
preparation and design of the substrate and the deposit
technique are factors that influence the performance of
the coating.

In addition to resorption, one must keep passive

dissolution in mind, which is a physiological process at a
buffer pH 7.4, and leads to localised precipitation of
amorphous apatite. However, at pH 5.2 (lactic acid),
this dissolution rate is multiplied by 40 [12]. Whereas we
know that active dissolution begins at pH 6.9 [37], all
recent studies have tried to prove the importance of the
environment on the evolution of the coatings. Findings
from animal experiments cannot be extrapolated to
human clinics, but these implants would be advanta-
geous in middle and poor bone quality, such as types 3
and 4, in post-extraction sites and after sinus bone
grafting, where they provided an earlier bone apposition
and primary stabilisation. But paradoxically, it is also in
this type of bone that less pronounced bone contact is
observed and where resorption of HA coating was more
pronounced. This work illustrates that HA coated
implants function the best adjacent to cortical bone. In
trabecular or cancellous bone would require coatings
with a very high crystallinity. In conclusion, the
behaviour of HA coatings on titanium implants placed
in the rabbit has been proven to be totally different
when a good bone contact is established between bone
and implant. Resorption of the coating appeared mainly
when bone quality is less dense. HA coatings showed a
higher percentage of bone contact than TPS coatings,
and the formation of bone was also accelerated in
presence of HA. Moreover, it was noted from an
ultrastructural observation that HA coating showed an
extremely high bonding strength with bone.

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