jba shekhar in vitro

8/18/2019 JBA Shekhar in Vitro

1/14

Article

In vitro biocompatibility of novelbiphasic calcium phosphate-mullitecomposites

Shekhar Nath, Sushma Kalmodia and Bikramjit Basu

Abstract

In designing new calcium phosphate (CaP)-based composites, the improvement in physical properties (strength, tough-ness) without compromising the biocompatibility aspect is essential. In a recent study, it has been demonstrated thatsignificant improvement in compressive strength as well as modest enhancement in toughness is achievable in biphasiccalcium phosphate (BCP)-based composites with mullite addition (up to 30 wt%). Herein, we report the results of thein vitro cell adhesion, cell proliferation, alkaline phosphatase (ALP) activity, and osteocalcin (OC) production for a seriesof BCP-mullite (up to 30 wt%) composites. Mouse fibroblast (L929) cell lines were used to examine in vitro cell adhesionand cell proliferation; while osteoblast-like (osteosarcoma, MG63) cells were used for in vitro osteoblastic function studyby ALP and OC expression. Much emphasis has been provided to discuss the cell viability and proliferation as well asosteoblastic differentiation marker on the investigated biocomposites in relation to the characteristics of the phaseassemblage. On the basis of various observations using multiple biochemical assays, it has been suggested that BCP-mullite composites would be a candidate material for orthopedic applications.

Keywords

BCP-mullite, composite, cell adhesion, MTT, ALP, osteocalcin

Introduction

In the search of ideal bioactive bone implant materials,

substantial efforts have been invested to develop

hydroxyapatite (HA) or CaP-based composites with

various reinforcements.1–4 As a synthetic analogue of

calcified tissues of vertebrate, HA is a candidate mate-

rial for bone implant applications. Also, bone matrix

can directly bind to HA, which is a necessary prerequi-

site for implant osseointegration.5 However, its low

strength and fracture toughness have reduced the field

of possible applications only to those, where the

implant will be subjected to very low stress.6–8

Another possible application of HA could be a coating

on metallic implants. However, recent clinical stud-

ies9,10 revealed that HA coating did not produce any

benefit, when implanted for long period. Keeping these

information in mind, it is therefore needed to improve

the performance of CaP (HA, tricalcium phosphate –

TCP]), which could be used as bulk. In search of such

materials, researchers used mixture of HA-TCP mate-

rials along with second phase reinforcement to develop

composites with desirable properties. To this end, the

physical property enhancement without any compro-

mise on biocompatibility aspect necessitates the use of

optimal amount of reinforcement as well as tailoring

the processing parameters.

For this purpose, in the present study, different

amounts (10–30 wt%) of mullite (3Al2O3.2SiO2) were

mixed with HA and the powder mixtures were sintered

under various optimal conditions. Mullite is a solid

solution of alumina (Al2

O3

) and silica (SiO2

). Mullite

is chemically inert and mainly used as refractory mate-

rial in high temperature furnaces.11 It has lower density

Journal of Biomaterials Applications

27(5) 497–509

! The Author(s) 2011

Reprints and permissions:

sagepub.co.uk/journalsPermissions.nav

DOI: 10.1177/0885328211412206

jba.sagepub.com

Laboratory for Biomaterials, Department Materials Science and

Engineering, Indian Institute of Technology Kanpur, Kanpur 208016, UP,

India

Corresponding author:

Bikramjit Basu, Laboratory for Biomaterials, Department of Materials

Science and Engineering, Indian Institute of Technology Kanpur, Kanpur

208016, UP, India.

Email: [email protected]


2/14

(3.05 g/cc) than Al2O3 (3.95 g/cc) and ZrO2 (6.1g/

cc). It has good combination of structural properties,

like high hardness of 15 GPa, high elastic modulus of

240 GPa, and moderate fracture toughness of

3 M Pa m0.5.12 When mullite was mixed (up to

30 wt%) with HA, very good combinations of mechan-

ical properties were measured (E-modulus: 70 GPa,hardness: 4–5 GPa, and fracture toughness: 1.5 MPa

m0.5).13 This toughness value was 2.5 times higher

than that of pure monolithic HA (0.6MPa m0.5,

measured by SEVNB technique). The compressive

strength of the developed composites was 230MPa,

which is by far better than pure HA (50 MPa).13

From microstructural phase assemblage point of

view, these composites contain a mixture of HA and

TCP and therefore they are called as biphasic calcium

phosphate (BCP)-mullite composites.14–16 Hence, it is

likely that such composite could be an excellent alter-

native of HA, provided that the composite shows

required biocompatibility.

In this backdrop, an important part of research is to

characterize the in vitro properties of new biomaterials.

Despite various conflicts in results of in vitro and in vivo

tests, the in vitro assays are considered the primary bio-

compatibility screening tests for a wide variety of

implant materials.17 For example, the response of oste-

oblastic cells to a thin film of poorly crystalline calcium

phosphate apatite (PCA) crystals was examined in vitro.

The osteoblasts were reported to exhibit high cellular

activity, such as adhesion, proliferation etc. In fact, the

cells were attached more rapidly to PCA thin film than

to reference dishes.18

In another study, HA was used asan additive for zirconia-alumina nanocomposite. The

addition of HA was reported to increase the biocom-

patibility of the nanocomposites significantly, as evi-

dent from the results of the in vitro tests using MG63

osteoblast-like cells.5 In some cases, the HA-based

composite materials showed better in vitro biocompat-

ibility than pure HA. For example, Boanini et al.19

studied the interaction of osteoblast-like cells on nano-

composite of HA with aspartic acid and glutamic acid.

Their results revealed that the nanocomposite of

HA possessed better cell proliferation, alkaline phos-

phatase (ALP) activity, and osteocalcin (OC) gene

expression than pure HA. Shu et al.20 studied the role

of HA on the differentiation and growth of MC3T3-E1

osteoblasts cells. Their results indicated that HA

enhanced osteoblast differentiation while suppressing

cell growth. In contrast, Licht et al.21 reported that

due to the phagocytosis of HA particles, osteoblast

exhibited reduced cell growth and ALP activity.

In another study, Ogata et al.22 compared the osteo-

blast response to HA with HA/soluble CaP (SCaP)

composites. Their results revealed that HA/CaP

showed greater ability in osteogenesis than HA by

increasing collagen synthesis and calcification of the

extracellular matrix.

Reviewing literature, it is evident that HA, TCP, and

other CaP phases are biocompatible. However, there is

no published data available on the biocompatibility of

mullite ceramics. Therefore, the biocompatibility of BCP-mullite composite materials is needed to be exam-

ined before any biomedical application can be pro-

posed. In this paper, we report the cell adhesion,

proliferation, and differentiation behavior using

mouse fibroblast (L929) as well as MG63 cell lines.

The sintered HA was used as baseline material in all

the in vitro experiments.

Materials and experimental procedures

Synthesis of materialsHA powder was synthesized in-house using commer-

cially available chemicals, such as calcium oxide

(CaO) and phosphoric acid (H3PO4), following a well-

established suspension–precipitation route.23,24 Phase

pure mullite (3Al2O3.2SiO2) powder was procured com-

mercially (KCM Corporation, Japan). As a first step of

sample preparation, the mixing of HA and mullite pow-

ders (10–30 wt% mullite) was carried out in a ball mill

for 16 h. Following this, the powder was subsequently

pressed to obtain pellets of 5 mm diameter. The sinter-

ing temperature was optimized based on the densifica-

tion and mechanical properties results. The sintering of the composite pellets was carried out at 1350C for 2 h,

while sintering of pure HA was performed at 1200C

for 2 h, both in conventional pressureless sintering fur-

nace. After the sintering, the diameter of the sample

was around 4 mm. The sintered composites are desig-

nated by their initial mullite content (BCP M means

BCP – x wt% mullite and likewise), irrespective of the

phases present in the sintered materials. Therefore,

throughout the text, such designation is followed for

the composites.

Material characterizationX-ray diffraction (Rich-Seifert, 2000 D) patterns were

acquired from the sintered ceramics to identify the dif-

ferent phases present. Based on the X-ray peak inten-

sity of the characteristic phases, the qualitative presence

of various phases is reported in this paper. Scanning

electron microscopy (SEM; model JSM-6330F,

Philips, The Netherlands) was performed on the pol-

ished and etched surface of the sintered composite to

investigate various microstructural features.

498 Journal of Biomaterials Applications 27(5)


3/14

Surface topography is an important parameter

which, influences the cell adhesion and, in general, the

biocompatibility property of materials. In order to

characterize the finer scale topography of the investi-

gated materials, smoothly polished and thermally

etched surfaces were observed under atomic force

microscope (AFM; Molecular Imaging, Pico-SPM I,USA) using ‘contact mode’. A Si3N4 cantilever with a

three-sided pyramidal single crystal Si3N4 tip with apex

angle of 20, a tip radius of curvature of 10 nm, and a

normal stiffness of 0.6 N/m were employed.

Cell culture experiment. L929 and MG63 cell linesobtained from CCMB, Hyderabad (India), and

ATCC (USA), respectably were preserved in an LN2container. Prior to seeding the cells on biomaterials

surfaces, the cells were revived. Following this, cells

were cultured in Dulbecco’s modified Eagle’s medium

(DMEM; Sigma Aldrich), supplemented with 10% fetal

bovine serum (Sigma Aldrich) and 1% penicillin/strep-

tomycin cocktail (Sigma Aldrich). The culture plate

with the cell lines (gelatin-coated) was incubated for

further proliferation and growth in a CO2 incubator

(Thermo, USA) operated under conditions of 5%

CO2, 90% humidity, and 37C temperature. The

medium was being replaced every 2 d interval and the

numbers of cells in confluent monolayer was approxi-

mately 5 105/mL, in 35-mm culture plate. The conflu-

ent monolayer was detached from the culture plate

using 0.50% trypsin and 0.20% EDTA solution

(Sigma Aldrich).

Cell adhesion test. As described in the previous sub-section (section ‘Cell culture experiment’), L929 cells

were cultured. The samples used for cell adhesion

experiment were pure HA, BCP10M, BCP20M,

BCP30M, and control glass disc. All samples were ster-

ilized in steam autoclave (121C, 15 lb pressure for

15 min) and, subsequently, the cells were seeded on

the samples at approximate density of 5 105/mL. In

a separate experiment, the cell seeding density of L929

was reduced to 1 105/mL and the effect of cell seeding

density on cell adhesion was studied. The seeded test

samples were incubated in a CO2

incubator with the

standard culture condition, 5% CO2, 37C tempera-

ture, and 90% humidity. The culture medium was

being aspirated after 2-d interval and fresh culture

medium was being added into each wells. After the

stipulated time period (1 and 3 d), the samples were

washed twice with phosphate buffer saline (PBS;

1PBS, pH 7.4) and then fixed by 2% glutaraldehyde

in PBS. The cells, adhered on the material surfaces,

were dehydrated using a series of ethanol solutions

(30%, 50%, 70%, 95%, 100%) for 10 min twice and

then further dried using critical point drier (CPD;

Quramtech, UK). The dried samples were sputter-

coated (Vacuum Tech, Bangalore, India) with gold

and examined under SEM. For AFM observation, the

dried samples were directly used, without gold coating.

The experiment was repeated for at least three times

and the representative results are presented in the pre-sent article.

MTT assay

L929 and MG63 cells were cultured following the pre-

viously described cell culture method. The cell prolifer-

ation was investigated on pure HA, BCP-mullite

composites (4 mm diameter, 4 mm height), and control

glass disc. At first, autoclaved samples were placed in

the 4-well plate and then washed with PBS. Following

this, 5 104 cells/mL were seeded on each sample.

Subsequently, the culture plate was incubated for 2 d

in the CO2 incubator. After the incubation period, the

medium was aspirated and samples were washed

twice with PBS. Then, 200mL of fresh DMEM was

added (without phenol red) into each well followed

by 10 mL reconstitute MTT (3(4,5-dimethylthiazol-2

yl)-2-5 diphenyltetrazolium bromide: SIGMA, USA,

Cat No.M5655; 5 mg/mL in DMEM culture medium

and serum) per 100 mL of DMEM was added in

each well and the plate was incubated for 6–8 h. In

the meantime, the culture plate was viewed under the

phase contrast microscope (Nikon, Eclipse 80i, Japan)

to check for the formation of purple formazane

crystal. After the incubation, samples were removedfrom the wells and placed in a new culture plate.

Thereafter, 200mL of dimethyl sulfoxide (DMSO)

was added into each well, including control. The opti-

cal density of the solution was measured at 540 nm

using ELISA automated microplate reader (Bio-Tek,

EL 800).

ALP activities

ALP is a widely recognized biochemical marker to

assess osteoblast activity on biomaterial substrate. It

is known that ALP plays a major role in skeletal min-

eralization. This enzyme is bound to the membrane of

osteoblasts and functions to enhance osteogenesis by

degrading pyrophosphates. For the ALP assay, the

autoclaved samples were placed in the 4-well plates

and MG63 cells were seeded approximately at a density

of 5 104 cells/mL. Three replicates of each composi-

tion were selected for this experiment. The culture pro-

tocol and conditions were similar to that described in

the earlier section. After 2 d of culture, ascorbic

acid and vitamin D3 were added to activate the

Nath et al. 499


4/14

phenotypic expression of differentiated MG63 cell line.

The concentration of 1,25-dihydroxy vitamin D3 solu-

tion has significant effect on the expression of differen-

tiated MG63 cells. In an important study, Robert and

Gideon25 reported that the concentration of 108 –107

M, 1,25-dihydroxy vitamin D3 inhibited growth and

elevated ALP as well as total cell protein. However,lower D3 concentrations (1010 and 109 M) reduced

ALP activity. Therefore, in the present study, the con-

centration of vitamin D3 was kept at 108 –107 M for

better ALP activity.

The ALP activity experiments were conducted after

3 rd and 7th d of culture. When the cells were lysed

followed by centrifuged. Therefore, the supernatant

was used as the ‘sample’ for the ALP activity. The sub-

sequent steps were followed as per the commercial kit

protocol (ALP, code no. 25904, Span Diagonstics Ltd.,

Surat, India) in a 96-well culture plate. At the last step

of the experiment, optical density was measured by

ELISA reader at 405nm.

Osteocalcin. It is known that OC, the most abundantnoncollageneous protein of the bone extracellular

matrix, is biologically synthesized by osteoblast-like

cells. In the present study, MG63 cells were cultured

for the OC assay, following the method described in

an earlier section. After subconfluent, cells were trypsi-

nized and seeded on sterilized pure HA and BCP-mul-

lite (all samples were of 4 mm diameter) in a 4 well

culture plate at a density of 5 104 cells/mL. After

day 2 and 5, culture medium was replaced by differen-tiating medium that contain vitamin D3 (108 M final

concentration), ascorbic acid (50mg/mL final), and

glycerol phosphate (108 mg/mL) to enhance osteogenic

activity. OC production was tested at 7th d of cul-

ture and the sample preparation was similar to ALP

assay. In brief, three strips of coated wells were taken

from the strips supplied with the OC kit (HOST-

EASIA KAP1381, Biosource Europe S.A., Belgium).

Thereafter, 25mL of each calibrator (human serum

with protease inhibitors and benzamidin: supplied

with kit), control (human serum with protease inhibi-

tors, benzamidin and thymol: supplied with OC kit)

and samples (Pure HA, BCP10M, BCP20M, and

BCP30M) were transferred into the appropriate wells

and subsequently 100mL of anti-OST-HRP conjugate

was added to all the wells, followed by 2 h shaking.

Following this, the liquid from each well was aspirated

and washed by 400 mL of wash solution. Then, the solu-

tion from each well was aspirated and 100 mL of chro-

mogenic solution was added into each well. Thereafter,

the well plate was incubated in a shaker at room tem-

perature. Finally, the optical density was measured at

450 nm.

Results

Phase assemblages and microstructure. In Table 1,the quantitative analysis of XRD results, obtained from

the polished surface of the investigated materials, are

presented. Also, a typical microstructure of BCP30M

composite is provided in Figure 1. The larger grainmatrix was identified as TCP/HA phase and the grain

boundary phases were mullite, alumina, CaO, and geh-

lenite. More details of the microstructural characteriza-

tion can be found elsewhere.15,16 From Table 1 it

should be clear that all the HA-based composites

after sintering predominantly contained TCP. In con-

trast, single-phase HA without any sign of dissociation

to any TCP polymorph could be obtained as baseline

HA. The addition of mullite as well as higher sintering

temperature, in combination, caused dissociation of

HA to TCP phases and it is clear that all the mullite-

containing samples were essentially BCP composites.

In Fig. 2, AFM images show some representative

surface topography (three-dimensional view). Figure

2(a) shows the equiaxed grains of HA of variable

sizes with good grain boundary adhesion. Figure 2(b)

shows the AFM image captured from BCP10M sample,

sintered at 1350C for 2h. Here, the finer grains of

reaction products were present at the boundary region

of larger HA/TCP grains. Similarly, Figure 2(c) shows

a coarse CaP-grain surrounded by the reaction prod-

ucts, that is, calcium–alumino silicate phase. Figure

2(d) presents a high magnification AFM image of

BCP30M sample and the grain boundary area was

almost covered by the reaction products. In all theinvestigated samples, the grain boundary phases

appeared to be well bonded with the matrix phase,

that is, HA/TCP. Also, the crystals of grain boundary

phases were much smaller (less than 300 nm) compared

to that of the matrix phase.

Cell adhesion. Figure 3 shows the cell adhesion behav-ior of BCP20M (Figure 3(a) and (b)), BCP30M

(Figure 3(c) and (d)), and control (Figure 3(e) and

(f)) samples, after 1 d of culture. Here, the cell seeding

density was 5 105 cells/mL. In all cases, cell division,

proliferation and cell–cell contacts were evident.

Clearly, as far as the cell adhesion was concerned, var-

ious composites, despite compositional difference/phase

assemblage, exhibited comparable behavior. Also, the

cell adhesion for investigated ceramics was comparable

with that of control sample. Beside major observation

of cell spreading, a transient signal of cell migration

was also observed through structural changes in the

form of cell protrusion.

More detailed observations of cell–material interac-

tion are made with SEM and AFM. Figure 4 shows the

results of similar cell adhesion experiments, but the



5/14

culture time was more (3 d) and the cell seeding density

was approximately 1 105/mL. SEM images reveal

adhesion of L929 cells on BCP10M (Figure 4(c) and

(d)), BCP20M (Figure 4(e) and (f)), BCP30M

(Figure 4(g) and (h)) samples. The cell adhesion on

baseline HA sample can be seen in Figure 4(a) and

(b). From the images, some interesting observations

can be made, for example, Figure 4(e) reveals how the

cells were connected to each other by filapodia exten-

sion on BCP20M sample. In this case, some observable

differences could be found, when the cell seeding den-

sity was lower. The cells were larger in size with an

approximate dimension of 40–50mm. The major differ-

ence could be found in cell morphology. In case of cell

density being 1 105/mL, most of the cells were spread

on the surface and intended to enhance the cell-material

contacts. The filopodia of the cells were extended over

longer distance than that was observed in the previous

case (when cell seeding density was 5 105/mL) on the

materials surface. Figure 4(i) shows an AFM image of a

filapodia attached on pure HA sample. The branching

of filapodia is clearly visible. This type of branching

provided good anchorage with the material surface.

MTT assay. It is known that MTT reagent directlyreacts with the mitochondria (mitochondrial dehydro-

genase) of living cells. Therefore, the reduction of MTT

will be more if more numbers of metabolically active

cells are present. In this context, MTT is widely

regarded as one of the quantitative assays to determine

the cytotoxicity of the materials, detecting the cell via-

bility on the sample surface. The measured optical den-

sity, as recorded with ELISA plate reader, is directlyproportional to the number of viable cells in the culture

medium.

Figures 5(a) and (b) plot the MTT assay results

obtained using MG63 and L929 cells, respectively. In

both the plots, the results are compared in reference to

pure HA. Figure 5(a) shows that in all the mullite-con-

taining composites, the numbers of metabolically active

cells were comparable to pure HA. BCP10M, BCP20M,

and BCP30M possessed 103%, 113%, and 101% cell

viability, respectively, compared to pure HA. Similar to

these results, Figure 5(b) shows the result of MTT assay

using L929 cells. Here also, the result obtained with HA

was considered as baseline observation (control).

Again, the MTT reduction rate data revealed that

both BCP10M and BCP20M, in average, possessed

similar or slightly better cell viability than pure HA.

Owing to the fact that error bars overlap among vari-

ous MTT datasets, no statistical significance could be

confirmed in terms of cell viability.

ALP activities. ALP activity is considered as a pheno-typic marker of differentiated cell. The ALP produced

by metabolically active MG63 cells, after 3 and 7 d of

Table 1. The starting powder composition and sintering conditions are mentioned as well as the sample designation of each sample

used in the present investigation.

Sample designation Starting composition

Pressureless sintering

conditions Phase assemblages (after sintering)

HA Pure HA 1200C, 2 h HA-ss

BCP10M HA with 10 wt% mullite 1350

C, 2 h a-TCP-ss, HA-ss, b-TCP-w, mullite-ww,CaO-ww

BCP20M HA with 20 wt% mullite 1350C, 2 h a-TCP-s, b-TCP-ss, HA-ww, mullite-w,

gehlenite-ww, CaO-ww, alumina-ww

BCP30M HA with 30 wt% mullite 1350C, 2 h b-TCP-ss, HA-ww, mullite-s, gehlenite-

ww, CaO-ww, alumina-ww

The phases present in the sintered ceramics are summarized based on the XRD peak intensities.

Ss: very strong, s: strong, ww: very weak, w: weak.

Figure 1. Representative microstructure of thermally etched

BCP30M ceramic. Note the uniform presence of grain boundary

phases around each grain.

Nath et al. 501


6/14

experiments, is quantitatively plotted in Figure 6. After

3 d of culture, the maximum ALP activity was mea-

sured with pure HA and for composites (10–30 wt%

mullite) the ALP activity was very near to pure HAp.

After 7 d of culture, the ALP expression of the cultured

cells was significantly higher in all mullite-containing

composite with respect to pure HA. However, almost

no difference in terms of ALP expression was found

among various mullite-containing composites.

Osteocalcin. OC is a later stage marker of bone celldifferentiation. Declercq et al.26 described calcification

as a predictor of bone mineralization capacity of bio-

materials in osteoblastic cell cultures. It is normally

accepted that the OC production measurement is

important to substantiate the use of investigated cera-

mics as bone replacement materials. As part of the pre-

sent study, the results of the OC assay, are plotted in

Figure 7. The OC production after 7 d of culture on

control, pure HA, and BCP10M showed significantly

lower values in comparison with BCP20M and

BCP30M samples. However, no difference in OC pro-

duction could be noticed between BCP20M and

BCP30M samples.

Figure 2. AFM images showing the surface topography of the various investigated materials: (a) pure HA sintered at 1200C for 2h

and the composites: (b) BCP10M, (c) BCP20M, and (d) BCP30M samples, all sintered at 1350C for 2 h. GB represents ‘grain

boundary.’



7/14

Figure 3. Selected SEM images illustrating the adhesion of L 929 cells on various material surfaces – BCP20M (a, b), BCP30 M (c, d),

and negative control sample (e, f) after in vitro culture for 1 d.

Nath et al. 503


8/14

Discussion

This study demonstrates that the biological response of

BCP (HA and TCP) ceramics containing 20 and

30 wt% mullite is comparable (as per cell adhesion,

MTT) or better (ALP and OC results) than pure HA.

It is indeed an important result that the presence of

mullite did not affect the cellular viability property on

BCP-mullite composites, when compared to sintered

HA monolith. In the following, results will be inter-

preted in terms of the substrate compositional differ-

ences or phase assemblage. Such interpretation will be

helpful to realize the following aspects: (a) In vitro cyto-

compatibility property on the basis of microstructural

phase assemblages; (b) prediction of osteoconduction

property based on biocompatibility with MG 63 cell

line; and (c) bone-cell functionality and phenotypic

marker expression based on ALP and OC expression

assay.

Materials and microstructure effect oncytocompatibility. In recent years, the mixture of HAand TCP ceramics has been reported to be better than

the monolithic forms of the either (HA or TCP) due to

the controlled resorbability of BCP ceramics. In an

interesting work, Arinzeh et al.27 studied the bone for-

mation capabilities of BCP ceramics using human

Figure 4. SEM images of L929 cells adhered on pure HA (a, b), BCP10M (c, d), BCP20M (e, f), and BCP30M samples (g, h). AFM

image revealing extensive filopodia extension on pure HA sample is also shown (i). Results were obtained after 3 d of culture. The

seeded cell density was 1105 cells/mL.



9/14

mesenchymal stem cells. The results revealed that HA :

TCP ratio of 20:80 was the optimum combination for

better bone formation, whereas pure HA and TCP

phase had much less effect on bone formation. A

closer look at Table 1 further reveals that while a-

TCP dominated in BCP10M material, b-TCP was the

major phase in BCP20M and BCP30M composites.

The additional presence of CaO, Al2O3, and gehlenite,

all in minor amounts (much weaker X-ray peak inten-

sity), was also noticed. From the examples mentioned

in introduction part, it is quite clear that pure HA and

TCP are undoubtedly biocompatible material.

However, the presence of other calcium-alumino-sili-

cate phases may also have an influence on the biocom-

patibility of this material. In some earlier reported

results,28,29 it was mentioned that the biocompatibility

of the materials mainly depended on the leaching of

ions. The adhered cells mainly proliferate and differen-

tiate due to the activation of surface ions. In the present

case, the possible leaching ions would be Ca, P, Si, and

Al. The effect of Ca and P ions need not be discussed

here, as these are already known marker for the bio-

mineralization.30 The effects of other ions, therefore,

need to be discussed with cited literature. For example,

silicon-containing materials were already investigated

as silicon-substituted HA, Si3N4, etc. Thain et al.31

(a)130

120

110

100

90

80

70

60

50BCP10M

M e t a b o l i c a l l y a c t i v e c

e l l s ( % P

u r e H A )

BCP20M BCP30M

(b)130

120

110

100

90

80

70

60

50BCP10M

M e t a b o l i c a l l y a c t i v e c e l l s ( % P

u r e H A )

BCP20MSamples

BCP30M

Figure 5. MTT assay results showing the relative number of

metabolically active (a) MG63 cells (b) L929 adhered on pure HA,

BCP10M, BCP20M, and BCP30M samples. In (a) and (b), theresults were compared with pure HA. The results are shown

after 2 d of culture.

3.00

2.50

2.00

1.50

1.00

0.50

0.00

Pure HA BCP10M BCP20MSamples

Days of culture

M e a n o p t i c a l d e n s i t y ( 4

0 5 n m )

3.007.00

BCP30M

Figure 6. ALP activity of MG63 cells on pure HA, BCP10M,

BCP20M, and BCP30M ceramics, after 3 and 7 d of culture in

osteogenic medium.

60

50

40

30

20

10

0Control Pure HA BCP10M

Samples

O s t e o c a l c i n ( n g / m L )

BCP20M BCP30M

Figure 7. OC expression of cultured MG63 cells on various

ceramics samples after 7 d of culture. The results were compared

with control solution supplied with the kit.

Nath et al. 505


10/14

described the in vitro biocompatibility of silicon-substi-

tuted HA thin film and reported that the presence of Si

did not induce any toxic effect. In another study, Kue

et al.32 showed enhanced cell proliferation and OC pro-

duction by human osteoblast-like MG63 cells on silicon

nitride ceramic discs. Their results revealed that Si3N4

was a nontoxic biocompatible ceramic, which could beused as a potential biomaterial. This proves that the

presence of Si in the composite, should not ide-

ally degrade the biocompatibility. As far as the Al-

leaching is concerned, a study by Ku et al. on Ti-6Al-

4 V alloy suggested that the release kinetics of Al-ions

could play a major role in influencing the osteoblast

behavior.28

Figure 1 shows some residual porosity mainly at

intragranular and few at intergranular regions. In gen-

eral, porosity degrades mechanical properties, but

microporosity has specific advantages during the initial

period of implantation. It should be mentioned here

that Hornez et al.33 reported that both mesoporosity

(10–50mm) and microporosity (1–10 mm) in HA essen-

tially stimulated significant cell growth of MC3T3-E1

osteoblast cells. The cell viability and cell functions

were better for microporous samples. The presence of

micropores would therefore allow better anchorage

with bone-forming cells and thus improved the mechan-

ical attachment of these materials during initial period

of implantation.34

Another parameter that is important in the context

of cell adhesion was the surface roughness. In the pre-

sent case, any minor difference in cell viability could be

attributed to difference in surface roughness. The sub-grains are clearly visible in the AFM images, while

those are not distinct in SEM image (Figure 1). From

Figure 2, it should also be clear that the presence of

large fraction of grain boundary phase increased the

local surface roughness at the nanoscale and this

should enhance the cell adhesion property in mullite-

containing BCP composites.

In addition, efforts were made to study the adhesion

and expansion behavior of single cell in isolation and in

contact with other cells on individual material surface.

This is important to know how a single cell expands or

attaches to material surface. The initial attachment of

cells is mediated by integrins, which induces dramatic

cytoskeletal changes leading to cell spreading, develop-

ment of focal adhesion complexes, cell migration, etc.

In fact, these are morphological signs, which are usually

described as overall cell morphology changes when a

cell recognizes the adsorbed adhesive proteins (and

their conformation) on a given material surface. The

cell adhesion and expansion of single cell on various

material surfaces can be seen in Figure 4(b, d, f, h).

Among various mammalian cell types, the fibroblast

is one of the least-differentiated cell lines. When

fibroblast cells were seeded at lower density, it

showed higher motility on the material surface. One

single cell can migrate and construct cell–cell interac-

tion with a recognizable polarity of movement.35 AFM

images, presented in Figure 4(i), reveal the evidences of

cell migration via the extension of cell filopodium. On

the other hand, lamellipodia extends the movementtoward the direction of cell travel and at the same

time adhere on the materials surface.35 When the cell

density was low, the cell–cell interaction was also lower

and, hence, the cells interact more with material surface

by spreading themselves on the surface (Figure 4(e)).

The unidirectional long filopodia extension from the

adhered cells often formed well-developed ‘Y’ or star

junction, as can be seen in Figure 4(e). In such a situ-

ation, cells were suitable to attach to the substrate and

grow in size as a part of the cell cycle in presence of

some growth factors.35 In contrast, high cell density

inhibits the growth of normal cells. In a cell crowding

surface, cell growth is inhibited by cell-to-cell contact

and as a result, it reduces spreading. The above discus-

sion corroborates well with the difference in cell size

and cell spreading, when cell seeding density was

varied (Figures 3 and 4).

The attachment of substrate-dependent cells, such as

fibroblasts and osteoblast, to a substratum (e.g. bioma-

terial) is a synchronized process involving cytoskeleton

reorganization, cell spreading, and formation of focal

contact.36,37 The cell adhesion behavior of pure HA

and composite samples showed presence of cytoplasmic

extension, filopodia (Figure 4(a-h)). From Figure 4(a)–

(h), it can be said qualitatively that the filopodia wasmore visible in case of pure HA (Figure 4(a) and (b)),

BCP20M (Figure 4(e) and (f)), and BCP30M

(Figure 4(g) and (h)) samples. However, the filopodia

was less visible and the cells surface had minimum

adhesion on BCP10M sample (Figure 4(c) and (d)).

BCP20M and BCP30M samples mainly contained

b-TCP (Table 1), whereas BCP10M mainly contained

a-TCP. Therefore, the difference in cytocompatibility

behavior could be attributed to the difference in phase

assemblage.

Osteoconduction and biochemical markersof bone turnover. As mentioned earlier, the aspectsof bone cell differentiation and functionality could be

explained on the basis of ALP and OC assay results.

The synthesis of these biochemical markers of bone cell

increases with the increasing expression of osteoblasts

and decreases with the maturation of osteoblasts.38 It is

well known that OC and ALP are the phenotypic mar-

kers of the late and early stages of differentiation of

osteoblast-like cells, respectively. An increased specific

activity of ALP in a bone cell essentially reflects a shift

toward a more differentiated state.



11/14

Recalling ALP results after 7 d, it is clear that all the

mullite-containing composites had comparable expres-

sion of early stage osteoblast differentiation marker,

which was much higher than baseline HA. This

means that osteoblast-like cells were in a better func-

tionally differentiated state in contact with the BCP-

mullite composites. Furthermore, the extent of OCexpression in case of BCP20M and BCP30M was

much higher than both baseline HA and BCP10M.

The above observations therefore confirmed that

BCP20M composite could exhibit the best combination

of OC and ALP expression.

Our results need to be explained in the perspective of

earlier literature reports. In several earlier research

reports, it was mentioned that among CaP-based mate-

rials, b-TCP containing composites showed better differ-

entiation and in vivo bone formation/mineralization

capability. For example, Shiratori et al.39 studied the

bone-forming ability of b-TCP, when implanted in

bone defects of rat femur. Their results revealed that

b-TCP was an appropriate material for the treatment

of bone defects. In another study, Matsuno et al.40 com-

pared the in vitro (ALP) as well as in vivo behavior of b-

TCP/collagen sponge composite with only collagen

sponge (CS). In the in vitro experiments, human mesen-

chymal stem cell lines were used and for in vivo experi-

ment, the samples were placed under the back skin of

nude mice for a time period of up to 12 weeks. Their

results revealed that the composite containing b-TCP

showed better osteogenic properties and had ability to

promote bone formation. In an interesting study,

Arinzeh et al.27

tried to optimize the optimum HA/b-TCP ratio for better stem cell–induced bone formation.

For this purpose, they selected various ratios of HA : b-

TCP, that is, 100:1, 76:24, 63:37, 56:44, 20:80, and 1:100.

In both in vitro (OC) and in vivo experiments (mouse), it

was revealed that HA:b-TCP ratio of 20:80 was the best

combination for new bone formation in bone remodel-

ing process. It can be further noted that this combination

showed better results compared to monolithic form of

HA and b-TCP. In fact, all other combination of HA

and b-TCP showed improved results compared to pure

HA and b-TCP. At the molecular level, HA phase pro-

vided the required binding sites, whereas TCP, due to its

higher dissolution properties, increased the concentra-

tion of Ca and P locally. All these led to the cascading of

differential gene expression and positively increased the

cell differentiation.41,42

In the present case, Table 1 clearly shows that both

BCP20M and BCP30M predominantly contained b-

TCP phase, while BCP10M contained more a-TCP

phase. On the basis of the above information, it

should therefore be clear that in view of the dominant

presence of b-TCP, both BCP20M and BCP30M exhib-

ited better expression of osteblastic phenotypic marker

than BCP10M. The present investigation also recon-

firmed that single-phase HA had inferior expression

of differentiated bone cell than BCP microstructure

containing varying ratio of b-TCP. Additionally, it

can be stated that the additional presence of gehlenite,

alumina, or CaO did not have any inhibitory effect as

far as the osteoconduction property of BCP20M andBCP30M was concerned. In addition to the differences

in phase assemblage, the difference in surface roughness

at nanoscale due to uniform presence of grain bound-

ary phase could also explain the observed difference in

osteogenic property in the present case.

In summary, it is clearly understood from MTT data

and cell adhesion tests that HA and BCP-mullite sub-

strates supported comparable number of viable cells.

This, in combination with all the above observations,

suggests that BCP-mullite biocomposites should be

used as a substrate for bone-forming cells.

Conclusions

a. The principal finding of the present study is that the

pressureless sintered BCP-mullite composites favor-

ably supported cell adhesion of L929 cells in vitro.

Also, such observations were independent of mull-

ite content (up to 30 wt%). The morphology and

motility of adhered cells was dependent on cell seed-

ing density.

b. As far as the quantification of the metabolically

active cells is concerned (cell viability), MTT

assay results with fibroblast and osteoblast-likecells did not reveal any statistically significant dif-

ference in BCP-mullite composites in comparison

with pure HA. This confirmed that mullite and

additional presence of various phases (alumina,

CaO, gehlenite) did not have any toxic effect

in vitro and the presence of mullite did not degrade

cell viability with respect to pure HA.

c. The combination of ALP activity and OC expres-

sion results indicates that BCP-based composite

substrates with 20% or 30% mullite supported

superior osteoconduction than baseline monolithic

HA ceramic. The presence of predominant b-TCP

phase was found to be suitable for osteoblastic

function and phenotypic expression. The presence

of additional phases, like gehlenite or alumina, was

not found to have any inhibitory effect in vitro.

Based on all the in vitro analysis and observations,

in combination, BCP-biocomposites containing

20% or 30% mullite can be considered as suitable

substrate for bone cell adhesion and proliferation.

d. AFM study in combination with SEM analysis

revealed the uniform presence of grain boundary

reaction phases, which increased the surface

Nath et al. 507


12/14

roughness at the nanoscale. The increased surface

roughness property could explain better cell adhe-

sion/proliferation property of BCP-mullite compos-

ites, compared to single-phase HA.

Acknowledgements

The authors wish to thank Department of Biotechnology

(DBT), Government of India, for the financial help. Thanks

to Dr Lakshmi Nair of CCMB, Hyderabad, India, for pro-

viding us L929 cells. Thanks are also due to Dr S. Ganesh and

S. Mittal of BSBE Department, IIT Kanpur, India, for their

help during initial period of cell culture experiments.

References

1. Nath S and Basu B. Designing materials for hard tissue

replacements. J Kor Ceram Soc 2008; 45: 1–29.

2. Nath S, Bodhak S and Basu B. HDPE-Al2O3-HAp com-

posites for biomedical applications: processing and char-

acterization. J Biomed Mater Res B Appl Biomaterials

2009; 88B: 1–11.

3. Nath S, Tripathi R and Basu B. Understanding phase

stability, microstructure development and biocompatibil-

ity in calcium phosphate-titania composites, synthesized

from hydroxyapatite and titanium powder mix. Mater Sci

Eng C 2009; 29: 97–107.

4. Nath S, Kalmodia S and Basu B. Sintering, microstruc-

ture and in vitro properties of HA-10 wt % Ag compos-

ites. J Mat Sci Mater Med 2010; 21: 1273–1287.

5. Kong YM, Bae CJ, Lee SH, Kim HW and Kim HE.

Improvement in biocompatibility of ZrO2 –Al2O3 nano-

composite by addition of HA. Biomaterials 2005; 26:

509–517.6. Adolfson E, Henning PA and Hermannson L. Phase

analysis and thermal stability of hot isostatically pressed

zirconia-hydroxyapatite composites. J Am Ceram Soc

2000; 83: 2798–2802.

7. Clifford A and Hill R. Apatite-mullite glass ceramics.

J Non-crys Sol 1996; 196: 346–351.

8. Kim HW, Kong YM, Koh YH and Kim HE. Pressureless

sintering and mechanical and biological properties of

flour-hydroxyapatite composites with zirconia. J Am

Ceram Soc 2003; 86: 2019–2026.

9. Camazzola D, Hammond T, Gandhi R and Davey JR. A

randomized trial of hydroxyapatite-coated femoral stems

in total hip arthroplasty a 13-year follow-up. J Arthrop

2009; 24: 33–37.

10. Gandhi R, Davey JR and Mahomed NN.

Hydroxyapatite coated femoral stems in primary total

hip arthroplasty. J Arthrop 2009; 24: 38–42.

11. Schneider H, Okada K and Pask J. Mullite and mullite

ceramics. New York: Wiley, 1994.

12. Kanzaki S, Tabata H, Kumazawa T and Ohta S.

Sintering and mechanical properties of stoichiometric

mullite. J Am Ceram Soc 1984; 68: C6–C7.

13. Nath S, Dubey A and Basu B. Mechanical properties

evaluation of novel calcium phosphate-mullite compos-

ites. J Biomat Appl , doi: 10.1177/0885328210393292.

14. Daculsi G, Laboux O, Malard O and Weiss P. Current

state of the art of biphasic calcium phosphate biocera-

mics. J Mater Sci Mater Med 2003; 14: 195–200.

15. Nath S, Biswas K, Wang K, Bordia RK and Basu B.

Sintering, phase stability, and properties of calcium phos-

phate-mullite composites. J Am Ceram Soc 2010; 93:

1639–1649.

16. Nath S, Biswas K and Basu B. Phase stability and micro-structure development in hydroxyapatite-mullite system.

Scripta Mater 2008; 58: 1054–1057.

17. Schottler FJ. Tissue culture methods for determining bio-

compatibility. ASTM STP 1983; 810: 19–24.

18. Hong JY, Kim YJ, Lee H W, Lee WK, Ko JS and

Kim HM. Osteoblastic cell response to thin film

of poorly crystalline calcium phosphate apatite

formed at low temperatures. Biomaterials 2003; 24:

2977–2984.

19. Boanini E, Torricelli P, Gazzano M, Giardino R and

Bigi A. Nanocomposites of hydroxyapatite with aspartic

acid and glutamic acid and their interaction with osteo-

blast-like cells. Biomaterials 2006; 27: 4428–4433.20. Shu R, McMullen R, Baumann MJ and McCabe LR.

Hydroxyapatite accelerates differentiation and suppresses

growth of MC3T3-E1 osteoblasts. J Biomed Mater Res A

2003; 67: 1196–1204.

21. Licht BA, Gregoire M, Orly I and Menanteau J. Cellular

activity of osteoblasts in the presence of hydroxyapatite:

an in vitro experiment. Biomaterials 1991; 12: 752–756.

22. Ogata K, Imazato S, Ehara A, Ebisu S, Kinomoto Y,

Nakano T, et al. Comparison of osteoblast responses to

hydroxyapatite and hydroxyapatite/soluble calcium

phosphate composites. J Biomed Mater Res A 2005; 72:

127–135.

23. Rathje W. Zur Kentnis de Phosphate: I. Uber hydroxy-

apatite. Bodenk Pflernah 1939; 12: 121–128.24. Santos MH, Oliveira M, Souza LPF, Mansurand HS and

Vasconcelos WL. Synthesis control and characterization

of hydroxyapatite prepared by wet precipitation process.

Mater Res 2004; 7: 625–630.

25. Robert JM and Gideon AR. The Effect of 1,25(OH)zD3

on alkaline phosphatase in osteoblastic osteosarcoma

cell. J Bio Chem 1982; 257: 3362–3365.

26. Declercq HA, Verbeeck RMH, Ridder LIFJMD,

Schacht EH and Cornelissen MJ. Calcification as an indi-

cator of osteoinductive capacity of biomaterials in osteo-

blastic cell cultures. Biomaterials 2005; 26: 4964–4974.

27. Arinzeh TL, Tran T, Mcalary J and Daculsi G. A com-

parative study of biphasic calcium phosphate ceramics

for human mesenchymal stem-cell-induced bone forma-

tion. Biomaterials 2005; 26: 3631–3638.

28. Ku CH, Pioletti DP, Browne M and Gregson PJ. Effect

of different Ti–6Al–4V surface treatments on osteoblasts

behavior. Biomaterials 2002; 23: 1447–1454.

29. Es-Souni M and Brandies HF. Assessing the biocompat-

ibility of NiTi shape memory alloys used for medical

applications. Anal Bioanal Chem 2005; 381: 557–567.

30. Koshihara M, Masuyama M, Uehara M and Suzuki K.

Effect of dietary calcium: phosphorus ratio on bone min-

eralization and intestinal calcium absorption in ovariec-

tomized rats. BioFactors 2004; 22: 39–42.



13/14

31. Thian ES, Huang J, Best SM, Barber ZH, Brooks RA,

Rushton N, et al. The response of osteoblasts to nano-

crystalline silicon-substituted hydroxyapatite thin films.

Biomaterials 2006; 27: 2692–2698.

32. Kue R, Sohrabi A, Nagle D, Frondoza C and

Hungerford D. Enhanced proliferation and osteocalcin

production by human osteoblast-like MG63 cells on sili-

con nitride ceramic discs. Biomaterials 1999; 20:1195–1201.

33. Hornez JC, Chai F, Monchau F, Blanchemain N,

Descampsand M and Hildebrand HF. Biological and

physico-chemical assessment of hydroxyapatite (HA)

with different porosity. Biomol Eng 2007; 24: 505–509.

34. Annaz B, Hing KA, Kayser M, Buckland T and

Silvio LD. Porosity variation in hydroxyapatite and oste-

oblast morphology: a scanning electron microscopy

study. J Microscopy 2004; 215: 100–110.

35. Freshney RI. Biology of cultured cells, culture of animal

cells: a manual of basic technique, 5th ed. New York: John

Wiley & Sons, Inc, 2005, p.31–42.

36. Misra A, Lim RPZ, Wu Z and Thanabalu T. N-WASPplays a critical role in fibroblast adhesion and spreading.

Biochem Biophy Res Comm 2007; 364: 908–912.

37. West KA, Zhang H and Brown MC. The LD4 motif of

paxillin regulates cell spreading and motility through an

interaction with paxillin kinase linker (PKL). J Cell Biol

2001; 154: 161–176.

38. Chang X, Hou ZM, Yasuaki S, Tomoo T and Akira Y.

Quantitative study of alkaline phosphatase and osteocal-

cin in the process of new bone. West China J Stomat

2005; 23: 424–426.

39. Shiratori K, Matsuzaka K, Koike Y, Murakami S,

Shimono M and Inoue T. Bone formation in b-tricalciumphosphate-filled bone defects of the rat femur: morpho-

metric analysis and expression of bone related protein

mRNA. Biomed Res 2005; 26: 51–59.

40. Matsuno T, Hashimoto Y, Nakahara T, Kuremoto K,

Nakamura T and Satoh T. b-TCP/Collagen sponge

composite enhances the osteogenic differentiation of

mesenchymal stem cells. J Hard Tissue Bio 2005; 14:

189–191.

41. Piattelli A, Scarano A and Mangano C. Clinical and his-

tologic aspects of biphasic calcium phosphate ceramic

(BCP) used in connection with implant placement.

Biomaterials 1996; 17: 1767–1770.

42. Uoshima MH, Ishikawa I, Kinoshita A, Weng HT andOda S. Clinical and histological observation of replace-

ment of biphasic calcium phosphate by bone tissue in

monkeys. Int J Periodont Rest Dent 1995; 15: 204–213.

Nath et al. 509


14/14

jba shekhar in vitro

Documents