ORIGINAL PAPER

In vitro biocompatibility of chitosan/hyaluronic acid-containingcalcium phosphate bone cements

Saeed Hesaraki • Nader Nezafati

Received: 19 October 2013 / Accepted: 24 December 2013

� Springer-Verlag Berlin Heidelberg 2014

Abstract The need for bone repair has increased as the

population ages. In this research, calcium phosphate

cements, with and without chitosan (CS) and hyaluronic

acid (HA), were synthesized. The composition and mor-

phological properties of cements were evaluated by X-ray

diffraction and scanning electron microscopy. The acellular

in vitro bioactivity revealed that different apatite mor-

phologies were formed on the surfaces of cements after

soaking in simulated body fluid. The in vitro osteoblastic

cell biocompatibility of in situ forming cements was

evaluated and compared with those of conventional cal-

cium phosphate cements (CPCs). The viability and growth

rate of the cells were similar for all CPCs, but better

alkaline phosphatase activity was observed for CPC with

CS and HA. Calcium phosphate cements supported

attachment of osteoblastic cells on their surfaces. Spindle-

shaped osteoblasts with developed cytoplasmic membrane

were found on the surfaces of cement samples after 7 days

of culture. These results reveal the potential of the CPC–

CS/HA composites to be used in bone tissue engineering.

Keywords Calcium phosphate cement � Chitosan �Hyaluronic acid � Alkaline phosphatase � Osteoblast

Introduction

Calcium phosphates (Ca–P) constitute an important group

of compounds widely used as synthetic bone substitution

materials [1–3], in clinical applications due to their good

biocompatibility, osteoconduction, simulation of bone

minerals, and bone replacement capability [4, 5]. The Ca–P

minerals also provide a desirable substrate for cell attach-

ment and expression of osteoblast phenotype [6, 7]. Cal-

cium phosphate cement (CPC) can be molded or injected to

form a scaffold in situ that properly adapts to the shape of

complicated bone defects [8]. Injectable scaffolds are

appropriate for cell delivery because they can lessen the

time of surgical operation, decrease postoperative pain and

scar size, make rapid recovery and reduce cost [9]. A

typical CPC was first reported by Brown and Chow [10].

This CPC powder could be mixed with aqueous liquid to

form a paste that would set in situ and form hydroxyapatite

as the main constituent part of the mineral phase of bone

[11]. Since then, several other calcium phosphate cements

and injectable cements have been developed [12–15]. On

the other hand, it is necessary to develop CPC with os-

teopromotive or osteoinductive factors to improve its bio-

logical performance [16, 17].

Hyaluronic acid (HA) is an anionic glycosaminoglycan

distributed widely through connective, epithelial and neural

tissues. In the extracellular matrix, hyaluronate can par-

ticipate as a hydrated network and operate as an organizing

core to connect complex intercellular aggregates [18, 19]

that assists significantly in cell proliferation and migration

[20]. Several applications of HA in tissue engineering have

been reported [21–23].

Chitosan (CS) has different biological properties that

make it attractive to medical applications including: bio-

degradability, lack of toxicity, antifungal effects, acceler-

ation of wound healing and stimulation of immune system

[24].

It has been reported that its handling, rheological

properties and antibacterial effect of CPC based on trical-

cium phosphate, tetracalcium phosphate (TTCP) and

S. Hesaraki � N. Nezafati (&)

Department of Nanotechnology and Advanced Materials,

Materials and Energy Research Center, Alborz, Iran

e-mail: n.nezafati@merc.ac.ir

123

Bioprocess Biosyst Eng

DOI 10.1007/s00449-013-1122-0

dicalcium phosphate (DCP) can get better if HA and/or

sodium hyaluronate and CS, as gel-forming polymers are

used [14, 25–27]. For example, Xu and colleagues have

shown that CS increases the flexural strength of a CS–CPC

composite composed of tetracalcium phosphate–dicalcium

phosphate anhydrous considerably, and the highest value

was reached when 15–20 wt% CS was incorporated into

the CPC, although Zhang reported that optimum results

also occurred when CPC–CS was synergistically combined

with Vicryl fibers or alginate microbeads [28]. Kai et al.

have investigated that the injectability, setting time and

mechanical strength of CPC could be significantly

improved by using 0.6 %wt of sodium hyaluronate (NaHA)

when NaHA, as the liquid phase, was introduced into the

system [29]. Hence, proper injectability along with bio-

compatibility can have good prospects for medical

application.

Despite the above-mentioned studies, it is worth to

mentioning that there is little information available con-

cerning the biological properties and evaluation of cell

behavior of CS/HA-based calcium phosphate cements for

bone composite cement. Thus, in the present study, after

the composition and morphological properties of cements

were assessed, apatite formation, cytocompatibility and

cell responses of the injectable CS/HA–CPC composites to

act as a bone substitute and its potential for bone tissue

engineering were investigated in vitro.

Experimental procedures

Sample preparation

The following starting materials were purchased and uti-

lized without further purification in this study: calcium

carbonate (CaCO3, Merck 2076, Germany), DCP anhy-

drous (Merck 2144, Germany), DCP dehydrate (DCPD,

Merck 172/06, Germany), high molecular weight HA

(1750 kDa; Jinan Haohua Industry Co., China), chitosan

(Chitotech Company, Iran) and acetic acid (Merck 100062,

Germany). The TTCP powder was synthesized by solid-

state reaction between equimolar amounts of DCP anhy-

drous and calcium carbonate, as described elsewhere [30].

Formulation of cements

The cement powder phase of the experimental groups was

a mixture of ground TTCP and DCPD (Brushite) powder in

a molar ratio of 1:1, and their liquid phases were prepared

by adding HA (CPC 1750) and CS (CPC–CS) at concen-

trations of 3 g ml-1 to a 1 % (v/v) solution of acetic acid.

Additionally, the pure CPC was also prepared with the

same powder phase and distilled water as a liquid phase to

compare its properties with those of the other two cements.

In each paste, the powder/liquid ratio (P/L) was selected in

a way that the best consistent paste can be obtained. This

proportion was 3/5 for CPC, 3/1 for CPC 1750 and

2/8 g ml-1 for CPC–CS. The powder phase was mixed

with the liquid, and the obtained paste was molded.

Characterization and measurement

Particle size distributions of synthesized powders of TTCP

and DCPD as a main part of cements were examined by

laser particle size analyzer (LPSA) via a Fritsch Particle

Size ‘Analysette 22’ instrument in the acetone–alcohol

mode.

Phase analysis of cement powder was performed by a

Philips PW3710 diffractometer. This instrument worked

with voltage and current settings of 40 kV and 30 mA,

respectively, and utilized Cu–Ka radiation (1.54 A). For

qualitative analysis, X-ray diffraction (XRD) diagrams

were recorded in the interval 10o B 2h B 50o.

The microstructure of the gold-coated fractured surfaces

of set cements was observed using SEM (Stereoscan S-360

Cambridge).

Acellular in vitro bioactivity

The in vitro surface reactivity of calcium phosphate-based

composites was accomplished after the samples were

soaked in simulated body fluid (SBF) at solid (S) to liquid

(L) loading of 1 g/100 ml and then kept at 37 �C for

14 days. The SBF solution was prepared according to the

procedure described by Kokubo et al. [31] by dissolving

NaCl 8.035 g/l, KCl 0.225 g/l, K2HPO4�3H2O 0.231 g/l,

MgCl2�6H2O 0.311 g/l, CaCl2 0.292 g/l, NaHCO3 0.355 g/

l and Na2SO3 0.072 g/l into distilled water, and buffered at

pH = 7.25 with 6.118 g/l tris–hydroxymethyl amino-

methane and 1 N HCl solution at 37C. The SBF solution

was chosen because of its characteristic of being highly

supersaturated with respect to apatite. The SBF solution is

so far the best for in vitro measurement of apatite-forming

ability in implant materials [32].

The surface morphology and microstructure of the

samples were evaluated using SEM after soaking in the

SBF solution for 14 days.

In vitro assay

The cement specimens (CPC, CPC–CS and CPC 1750)

were selected for clinical applications because of their

ability in setting and also due to diverse usage of such

materials in reconstruction of bone injuries. In this part of

the study, polystyrene pieces with dimensions similar to

that of cement specimens were used as control. It is

Bioprocess Biosyst Eng

123

necessary to apply biological tests for studying biocom-

patibility of cement materials. In this study, to evaluate the

biological properties of pastes, osteoblastic cells were

derived from newborn rat calvaria. These cells were pur-

chased from Royan Institute Research Center.

Sample preparation

The cylindrical cement pastes were formed in Teflon mold

(10 mm in diameter and 3 mm in height) for measuring the

level of cytotoxicity and cell morphology. To obtain

specimens with smooth surfaces, the molds filled with

cement pastes were packed with a glass plate. After the

cements hardened, the samples were sterilized with 70 %

ethanol and seeded with the cells [33]. The cultured cells

were continuously checked by a microscope to change the

cell medium if it was necessary.

Cell culture procedure

The cells were cultured in Dulbecco modified Eagle med-

ium (DMEM;Gibco-BRL, LifeTechnologies, GrandIsland,

NY) supplemented with 15 % fetal bovine serum

(FBS;Dainippon Pharmaceutical, Osaka, Japan) in a 5 %

CO2 atmosphere at 37˚C. The size of the cell flask was

50 mm. To store and freeze the cells, a suspension of cells

(2 9 104 cells for each vial) was poured into a 0.5 ml

medium including 90 % fetal bovine serum and 10 %

dimethylsulfoxide (DMSO). This suspension was conveyed

to cryotubes and frozen at -20 �C for 2 h. The tubes were

frozen at -70 �C and then transferred to a liquid nitrogen

tank after about 4 h. To subculture the cells, the tubes were

heated at 37 �C in a Benmary bath and the cells were

washed with culture medium after the ingredients of the

tube were melted. Finally, the resultant cell suspension was

cultured. The confluent cells were dissociated with trypsin

and subcultured to three passages which were used for

tests.

Cells proliferation assay and alkaline phosphatase (ALP)

activity

The osteoblastic cells (cultured for 1 week) were cultured

on the surfaces of cements (cell density: 2 9 104 cell/

sample) and incubated for 4 h. The specimen/cell samples

were placed in 24-well culture plates and left undisturbed

in an incubator for 3 h to allow the cells to attach to them

and then an additional 30 ml of culture medium was added

into each well. The cell/specimen constructs were cultured

in a humidified incubator at 37 �C with 95 % air and 5 %

CO2 for different periods. Note that every 3 days, the

medium was exchanged.

The proliferation of the osteoblastic cells was deter-

mined using the MTT (3-{4,5-dimethylthiazol-2yl}-2,5-

diphenyl-2H-tetrazolium bromide) assay. For this purpose,

at the end of each evaluating period, the medium was

removed and 2 ml of MTT solution was added to each

well. Following incubation at 37 �C for 4 h in a fully

humidified atmosphere at 5 % CO2 in air, MTT was taken

up by active cells and reduced in the mitochondria to

insoluble purple formazon granules. Subsequently, the

medium was discarded and the precipitated formazon was

dissolved in dimethylsulfoxide, DMSO, and optical density

(OD) of the solution was read using a microplate spectro-

photometer (BIO-TEK Elx 800, Highland park, USA) at a

wavelength of 570 nm. The recorded OD values were

converted into cell number by preparing a calibration curve

using known cell numbers.

The osteoblast activity was determined by measuring the

level of alkaline phosphatase enzyme. The unit of ALP is

defined as the amount of enzyme that hydrolyzes 1 lmol of

p-nitrophenyl phosphate to p-nitrophenol in a total reaction

volume of 1 ml in 1 min at 37 �C [34]. The cells were seeded

on the samples under the same culturing condition described

elsewhere and the level of ALP was ascertained on different

days. The osteoblast lysates were frozen and thawed three

times to disrupt the cell membranes. ALP activity was

determined at 410 nm using p-nitrophenyl phosphate in

diethanolamide buffer as chromogenic substrate.

To observe the morphologies of the cells attached onto

the surfaces of the specimens, after 7 days the culture

medium was removed, the cell-cultured specimens were

rinsed with phosphate-buffered saline (PBS) twice and then

the cells were fixed with 500 ml/well of 2.5 % glutaral-

dehyde solution. After 30 min, they were rinsed again and

kept in PBS at 4 �C. The specimens were then fixed with

1 % osmium tetroxide. After cell fixation, the specimens

were dehydrated in ethanol solutions of varying concen-

trations (30, 50, 70, 90, and 100 %) for about 20 min at

each concentration. The specimens were then dried in air,

coated with gold and analyzed by SEM.

Results

Particle size distribution and specific surface area

According to the curve of particle size distribution, the

mean particle diameter of TTCP and brushite powders was

determined to be 12/28 and 6/13 lm, respectively. It was

also determined that 90 % of TTCP particles had a size less

than 30 lm (Fig. 1a), whereas 90 % of DCPD particles had

a size \19 lm (Fig. 1b). Therefore, the mean particle

diameter of TTCP particles was generally two times more

than brushite particles.

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Phase analysis of the specimens

Figure 2 depicts the compositional variations in the CPC

samples after being set for 24 h in an incubator. In all three

groups of set cements, the formation of the apatite phase,

which is the main product of the setting reaction, could be

detected along with the TTCP and DCPD phases as

reactants.

SEM observation of samples before and after soaking

in SBF solution

The microstructures of CPC, CPC–CS and CPC 1750

samples after 24 h incubation at 37 C are shown in Fig. 3.

According to Fig. 3a, two different morphologies of cal-

cium phosphate reactant particles were observed for the

CPC sample: first, coarse particles with dimension of

nearly 15 lm; second, finer particles (\5 lm) which were

situated on the surface of larger ones. In addition, the mi-

cropores (\3 lm) were produced by removing the liquid

phase from the cement structure. The microstructure of

calcium phosphate particles in CPC–CS (Fig. 3b) and CPC

1750 (Fig. 3c) was completely different from the CPC

sample. It seemed that reactive components in these

cements were covered with bindery phase, and the surface

of particles was smoother in these samples. The tiny holes

on the surface of CPC–CS and CPC 1750 were probably

generated by the air-trapped bubbles.

Scanning electron microscopy images of samples, after

14 days of soaking, showed that tiny spherical apatite

particles were formed on the surface of CPC (Fig. 4a),

while the morphology was changed for CPC–CS (Fig. 4b)

and CPC 1750 (Fig. 4c). For these samples, the particles

were connected by entangled needle-like apatite crystals. It

seems that this kind of morphology is nearly more

detectable for CPC 1750 than CPC–CS.

Fig. 1 Particle size distribution curve of TTCP and brushite powders by LPSA analyzer

Fig. 2 XRD pattern of 1 CPC,

2 CPC–CS and 3 CPC 1750

specimens after being set for

24 h in an incubator

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Cell viability

Figure 5 shows the function of calcium phosphate cement

with and without polysaccharide-based polymers on the

viability of osteoblastic cells measured by MTT assay on

days 1, 7 and 14. The principle of MTT assay is based on

the conversion of MTT into formazan crystals by living

cells, which determines mitochondrial activity. For most

viable cells, mitochondrial activity is constant and thereby

an increase or decrease in the number of viable cells is

linearly related to mitochondrial activity. Since for most

cell populations the total mitochondrial activity is related

to the number of viable cells, this assay is broadly used to

measure the in vitro cytotoxic effects of cement constitu-

ents on cell lines or primary patient cells [35]. The number

of viable cells on the surfaces was equal for all samples on

day 1. After some time, the number of cells was consid-

erably increased. This issue confirmed the lack of cell

Fig. 3 SEM micrographs of

a CPC, b CPC–CS and c CPC

1750 samples after 24 h

incubation at 37 �C at

magnification of 91500

Fig. 4 SEM micrographs of

a CPC, b CPC–CS and c CPC

1750 after 14 days of soaking in

SBF solution

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123

toxicity in the medium and exhibited an initial biocom-

patibility of the samples. On the other hand, the growth rate

of cells was the same for all CPCs. The number of cells on

the surfaces of CPC–CS and CPC 1750 was also alike in

comparison with pure CPC samples. The growth rate of

cement samples was higher than that of the control. With

regard to surface morphology, it means that contrary to

control samples, the surface of cement samples was porous

(according to SEM images), producing a proper situation

for cell attachment.

Alkaline phosphatase activity

The biocompatibility of the samples was assessed through

in vitro cell culture experiments. In this way, focus has

been on alkaline phosphate activity (ALP activity) of

osteoblastic cells, because one of the phenotypic markers

for osteoblast proliferation and differentiation is alkaline

phosphatase expression. The osteoblastic cells on the

samples were assayed for retention of their osteoblast-like

phenotype and the results are shown in Fig. 6. On day 1,

alkaline phosphate activity was equal for all samples.

This value was higher for cement samples than control

after 7 days. This is because of calcium and phosphate

ions release from the cement matrix in the culture med-

ium that affects the enzymatic activity of the cells. The

presence of calcium ions in the extracellular matrix

causes an increase in cell activity. However, as it can be

seen in the figure, ALP activity at 14 days was decreased

compared with that at 7 days. On the other hand, although

differences between normalized ALP values of various

cements were not statistically significant, the resultant

average values of HA samples were higher than those of

others.

Cell morphology on the surface of samples

Figure 7 depicts the morphology of osteoblastic cells on

the surfaces of cement samples after 7 days. The cyto-

plasmic membrane of cells spread widely and was attached

on the surface of three types of samples. This represented

high biocompatibility and showed that these surfaces could

support cells to become viable. The microporous structure

of cements provided such an appropriate attachment and

differentiation.

Discussion

In our experiments, various calcium phosphate cements

were made with and without CS and HA. The composition

and morphological properties of cements were also com-

pared to each other (Fig. 3).

Basically, a reaction between TTCP and DCPD as

reactants does not make a medium acidic or basic, but the

range of pH will be retained between 7 and 8. The rate of

solubility for TTCP and DCPD is different as this param-

eter is higher for TTCP than DCPD [36]. On the other

hand, solubility of TTCP causes increasing pH, because

this material hydrolyzes during the following reaction:

3Ca4 PO4ð Þ2þ3H2O! Ca10 PO4ð Þ6 OHð Þ2þ2Ca OHð Þ2 ð1Þ

The value of pH will increases due to the presence of

Ca(OH)2; this situation can inhibit reaction between reac-

tants and finally apatite formation. So, to equal the rate of

solubility of these materials, the size of TTCP particles

should be considered to be larger than DCPD particles. The

pH always remains in the range of 7–8 and apatite for-

mation will continue. As a result, particles size will affect

the chemical reaction of components. The process of apa-

tite formation, resulting from the reaction between reac-

tants, has been discussed by some researchers [37]. They

reported that the initial crystals precipitated on the surface

of DCPD particles produced apatite with complete stoi-

chiometric structure, but in the next step of the process they

were converted into calcium-deficient hydroxyapatite

Fig. 5 Proliferation of the osteoblastic cells on various specimens

Fig. 6 ALP activity of the osteoblastic cells cultured on the sample

for 14 days

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(CDHA). The presence of the initial apatite crystals was

due to the solution of TTCP particles, their hydrolysis to

stoichiometric apatite Eq. 1) and the growth of apatite

crystals. The growth of apatite crystals is because of the

acidic–basic reaction which is shown below:

3Ca4 PO4ð Þ2Oþ 6CaHPO4 ! 2Ca2 HPO4ð Þ6 PO4ð Þ5OH

þ H2O: ð2Þ

However, other researchers believe that the initial

crystals precipitated on the surface of DCPD particles are

non-stoichiometric apatite or CDHA. They describe that

this event may occur because of the dissolution of layers

related to DCPD crystals and formation of a supersatura-

tion situation in the medium around the particles [38]. The

dissolution rate of these particles is decreased by the DCPD

particles which have been covered by non-stoichiometric

apatite crystals. In addition, when the hydrolysis process of

TTCP particles goes on (see Eq. 2), the pH value increases.

Thus, these mechanisms influence hydroxyapatite compo-

sition. It means that in the middle stage of the process,

stoichiometric apatite is formed by hydrolysis of TTCP

and, subsequently, non-stoichiometric apatite crystals are

formed by reaction between the rest of the DCPD and

hydroxyapatite particles. In fact, according to this

approach, there are three stages during the precipitation

and growth processes: first, nucleation; second (middle

stage), formation of stoichiometric apatite crystals; and

last, formation of CDHA crystals. It is worth mentioning

that the structure of non-stoichiometric apatite is similar to

stoichiometric apatite. This can be confirmed by XRD

pattern. It has been reported that the carboxyl groups on

material surfaces are appropriate sites for the nucleation of

apatitic crystals [39]. HA contains COOH functional

groups, while CS is composed of NHAc groups and this

can describe why CPC 1750 has more detectable and

thicker apatite crystals than other samples. In other words,

the COOH groups can favor apatite precipitation and

growth over CPC 1750 surface, though these crystals are

thoroughly formed over the surfaces of other samples.

For CPC–CS, chitosan can produce some complexes

with calcium ions. Calcium phosphate-based composite

material was usually prepared as a bioactive layer to ini-

tiate the formation of hydroxyapatite layer, especially for

the CS/calcium phosphate composites. Furthermore, the

biologically related investigation of composite materials

proved that CS served only as a supporting material, while

the calcium phosphate phase played an important part in

inducing the nucleation of hydroxyapatite in the formation

of surface coating [40–42]. In this research, the mixture of

CPC and CS led to increasing pH and slow conversion of

reactants to hydroxyxapatite. According to Fig. 2, the XRD

patterns of three kinds of CPCs are similar to each other

and have no significant difference. In terms of the mor-

phology of apatite, needle-like apatite formation is a result

of a series of hydraulic reactions between reactants. An

Fig. 7 SEM micrograph of osteoblast cells grown on a calcium

phosphate cement (CPC), b calcium phosphate cement containing

hyaluronic acid (CPC–HA) and c calcium phosphate cement

containing chitosan (CPC–CS) for 7 days

Bioprocess Biosyst Eng

123

important regulating variable in the crystallization of nee-

dle-like crystals is nucleation. According to Fig. 4, this

event is true for CPS–CS and CPC 1750 samples. Basi-

cally, dissolution–precipitation mechanism caused this

kind of apatite morphology on the surface of samples.

According to the reports of other researches [43], the

microstructure morphology created in the simulated phys-

iological media is controlled by the solubility of the dif-

ferent phases, which generates changes in the surface

chemistry and the surface topography. The nucleation of a

calcium-deficient hydroxyapatite layer also takes place by

reaction of the SBF phosphate ions with the excess of

calcium ions released into SBF by the material. In fact, the

apatite will start nucleating and then grow spontaneously in

the SBF if the sample brings up a suitable nucleation site or

the degree of supersaturation exceeds a specific limit for

heterogeneous or homogeneous nucleation via the release

of constituent ions of apatite from the sample. Therefore,

formation of such rugged surfaces after 14 days of soaking

can be because of the above-mentioned mechanism [44].

A successful synthetic bone grafting system consists of

three factors: the extracellular matrix, diffusible growth

factors/proteins and viable cells [45]. Viable cells con-

tribute to the proliferation, differentiation or induction and

ultimate mineralization of bone tissue. The cells also pro-

vide a local source for soluble growth factors found in the

ECM [45]. In this research, to obtain an injectable paste for

bone repair, we prepared calcium phosphate and mixed

them with HA and CS. The cellular responses to bioma-

terials can be influenced by surface characteristics of the

biomaterials in vitro. Ideally, bioactive materials should

interact actively with cells and stimulate cell growth [46].

Its quality will directly affect cell growth, morphology,

proliferation and differentiation, because the cell attach-

ment stage is the initial stage of interaction between the

cells and the biomaterial [47]. Based on results, cells

adhered well on the surface of the three types of samples.

Biocompatibility is also a factor relevant to the response

of cells that are in contact with the biomaterial, and it has

been reported that the surface of biomaterials may affect

the behavior and morphology of cells cultured on their

surface [48]. SEM results for cell morphology confirmed

that cells attached and spread on the surfaces of three types

of samples. Moreover, cellular responses to biomaterials,

such as cell attachment, proliferation and differentiation,

depend not only on the surface morphology but also on the

chemical composition of the biomaterial [49], which plays

a main role in determining the cell–material interaction for

biomaterials by influencing the quantity of ions released

from the biomaterial [50]. For example, Huang et al.

reported that CS delivered as microparticles induced pro-

inflammatory responses in rat lungs. They mentioned that

sensitivity of cells was also likely to vary between different

cell lines and under different culture and assay conditions

[51]. In another investigation, SEM images published by

Fakhry et al. indicated that osteoblasts attached on CS

substrates at a higher rate [52]. Hsiao also proved that there

was a significant and specific affinity between CS and

osteoblasts. The adhesion force between CS and osteo-

blasts or osteoblastic cells was found to be particularly high

[53]. The results of Akmal et al., showed that articular

chondrocytes cultured in the presence of HA possessed a

significantly greater rate of DNA proliferation and extra-

cellular matrix production, compared with chondrocytes

cultured without HA [54].

According to the results, CPC–CS and CPC 1750 did not

adversely affect cell proliferation compared to traditional

CPC, which was approved by the Food and Drug Admin-

istration for craniofacial repairs [55, 56]. The incorporation

of HA and CS into the CPC composition had a positive

effect on the osteoblastic induction of cells in vitro. This

indicated that there was no significant difference between

pure CPC and CPC–CS/CPC 1750 composites in terms of

cell viability, cell attachment and osteoblast proliferation

and differentiation. According to Figs. 5 and 6, although

cell number of all samples was increased with culture time,

ALP activity at 14 days was decreased compared with that

at 7 days. This can be described as follows. The growth of

cell monolayers in vitro can be divided into three main

stages [57, 58]. In the early stage when the cell density over

the surface is low, the cells grow exponentially. Then the

growth rate decreases as a consequence of the fact that cells

come in contact and form colonies. In this part, cell pro-

liferation within each colony is mainly restricted to their

boundary. Hence, the colonies’ diameter grows linearly at a

constant velocity. Eventually, if the colonies fill the entire

domain, the growth rate decreases till it stops. Then, it is

said that the monolayer has grown to confluence. The

related control mechanism is called contact inhibition of

growth. So, basically, contact inhibition in confluent cell

cultures is currently defined as: (a) a dramatic decrease of

cell mobility and mitotic rate with increasing cell density;

(b) establishment of a stationary postconfluent state which

is insensitive to nutrient renewal. It is widely believed that

contact inhibition is caused by cell contact. But despite

extensive study, the current understanding of the mecha-

nism of contact inhibition is far from complete.

Conclusion

In this study, injectable CPC and CPC 1750 and CPC–CS

were prepared. It was concluded that the CPC–CS/CPC

1750 showed a suitable biocompatibility for its use as cell

culture composite for hard tissue regeneration. Spherical

and needle-like shapes of apatite were formed on the

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surfaces of cements when they were soaked in SBF solution

for 14 days. In vitro experiments revealed that cell viability

and growth rate of cells for CPC–CS and CPC 1750 samples

was nearly similar. Collectively, these data indicated that

CPC 1750 may give greater results concerning average

value of ALP compared to the CPC–CS and pure CPC in this

in vitro study. Overall, we suggest an acceptable applica-

bility of the CPC–CS/CPC 1750 samples as bone

substitutes.

Acknowledgments The authors would like to acknowledge the Iran

National Science Foundation (INSF) for the financial support of this

work through Grant No. 89001740.

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