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16.4 Nanocoatings and Surface Modiﬁ cations 539

For these studies, the coral was obtained from the Australian Great Barrier Reef

and contained micropores of 100 to 300 μ m size (Figure 16.6 ). The coral was

shaped in the form of a block, and treated with boiling water and 5% NaClO solu-

tion. A hydrothermal conversion was carried out in a Parr reactor (Parr Instrument

Company, USA) with a Teﬂ on liner at 250 ° C and 3.8 MPa pressure with excess

(NH

)

HPO

. A total conversion to HAp was achieved in this way.

Nanocrystalline ceramic coatings were produced using the sol – gel process. For

this, the precursor solution was formed using the previously reported method.

Coatings were formed using these solutions, followed by subsequent heat treat-

ments. Mechanical testing involved a standard, four - point bend test according to

ASTM C1161, to measure the ﬂ exural strength and ﬂ exural modulus of the natural

coral. Comparative compression and biaxial strength tests were also carried out.

Fracture surfaces were then viewed using SEM, which was performed on a LEO -

Supra55VP instrument. Samples were analyzed using X - ray diffraction ( XRD ;

Siemens D - 5000, Karlsruhe, Germany), with scans being carried out from 20.0 to

60.0 in 0.020 steps at a step time of 2.0 s. A combined thermogravimetric analysis

( TGA )/ differential thermal analysis ( DTA ) was performed using a TA Instruments

SDT 2960, at a heating rate of 10 ° C min

− 1

Characterization studies of the natural and converted corals using XRD, SEM,

DTA/TGA, nuclear magnetic resonance ( NMR ) and Raman spectroscopy have

been reported previously.

Figure 16.6 (a) Structural differences and morphology of (a) a

synthetic tricalcium phosphate and (b) a natural Australian

coral skeleton, showing pore size, distribution, and

interconnectivity.

540 16 Nanoceramics for Medical Applications

These results showed a large increase in all mechanical properties, speciﬁ cally

the compression strength, due to hydrothermal conversion and nanocoating

methods. The bioactivity was enhanced through the nanocrystalline formation,

due to the HAp nanocoating.

16.4.3

Surface Modiﬁ cations

The surfaces of nanostructured materials can be modiﬁ ed and functionalized with

different reagents, using a variety of physical, chemical, and/or biological methods.

An enhanced solubility or stability of nanosized materials in aqueous media, as

well as new material functions and properties, can be achieved via the surface

modiﬁ cation of nanostructured materials.

A variety of physical, chemical, and biological surface modiﬁ cations to increase

bioactivity or mechanical properties have been proposed and investigated by many

groups. Molecular coating, surface entrapment, and physical treating with plasma,

ozone, or ultraviolet ( UV ) light have emerged as the leading strategies for surface

modiﬁ cations of nanostructured materials using physical methods. Through

physical modiﬁ cations, a range of functional molecules and entities, varying

charges or active chemical groups can be introduced onto the surfaces of nanos-

tructured materials, leading to functionalization and activation of the surfaces

of materials.

Functional molecules may also be linked to the surfaces of nanostructured

materials via certain chemical reactions. Compared to certain physical methods,

a chemical modiﬁ cations can be used not only to activate the surfaces of nanos-

tructured materials to a greater extent, but also to offer stronger interactions

between the linking molecules and the material surfaces through stable chemical

bonds. A number of different chemical reagents and methods can be used for the

surface modiﬁ cation of nanosized materials.

The biological modiﬁ cation of nanoparticle surfaces is often necessary for nano-

particle functionality. By employing chemical or physical methods, biospeciﬁ c

molecules and devices can be incorporated into the nanoparticles, thereby offering

biospeciﬁ c sites for the further immobilization of ligands speciﬁ c to these mole-

cules. The immobilizations of speciﬁ c ligands can be performed through biologi-

cally speciﬁ c reactions, for example antibody – antigen and receptor – ligand [21] .

Different biomedical devices and applications require different properties and

functions of materials. Therefore, methods to modify nanostructured materials in

order to meet the needs of various biomedical systems will vary. A brief summary

of the basic methods and technologies used to modify nanostructured materials

for biomedical devices is presented below.

Today, highly porous scaffolds with an open structure represent the best candi-

dates for cancellous bone substitution. In addition to natural and ceramic materi-

als, many polymers have been proposed for medical applications. Each of these

presents different biological and mechanical properties, allowing a choice of the

correct polymer for the correct application. However, polymers usually present low

16.5 Simulated Body Fluids 541

elastic modulus values, creep resistance and chemical constituents, compared to

bone, and this is the major reason that limits their clinical use for hard - and soft -

tissue substitution.

One example was that reported by Peroglio and coworkers [11] , who produced

PCL - coated alumina scaffolds that were then characterized to validate the concept

of polymer – ceramic composites with an increased fracture resistance. The alumina

scaffolds were processed using a classical foam replication technique, and then

sintered to produce an open - porous structure with ∼ 70% porosity and a mean pore

size of 150 μ m. The polymer coating was obtained by inﬁ ltrating the scaffold with

either a PCL solution or PCL nanodispersion. An emulsion – diffusion technique,

using a nonionic surfactant, was used for the latter process. Subsequently, after

inﬁ ltration with PCL, and irrespective of which quantity or inﬁ ltration technique

was used, no change was seen in the Young ’ s modulus. However, this was to be

expected as the elastic modulus of PCL is negligible compared to that of alumina.

Nonetheless, the addition of PCL completely altered the mechanical behavior of

the scaffold during a four - point bending test, with a 10 – 20 vol% addition of PCL

to the alumina scaffold leading to seven - to 13 - fold increases in the apparent frac-

ture energy. A further examination of the material, using SEM, indicated that the

toughening was the result of the polymer ﬁ brils bridging the cracks.

In recent years, nanoparticle systems have attracted increasing attention for

use as potential drug - delivery systems. Despite the advantages of nanoparti-

cles – such as their small size, which allows them to penetrate small capillaries

and be taken up by cells – a number of problems, including a relatively short

blood circulation time, have limited their clinical application. The efﬁ ciency and

targeting ability of a nanoparticle drug - delivery system are often hampered by

the rapid recognition of the carrier system by the body. As the main concern for

nanoparticle drug carriers is a long circulation time in the blood, numerous

approaches to the design and engineering of long - circulating - time carriers have

been investigated. Among these, the surface modiﬁ cation of nanoparticles with

a range of nonionic surfactant or polymeric macromolecules has proved to be

the most successful for maintaining nanoparticle presence in the blood for pro-

longed periods [41] . Suitable and effective modiﬁ cations of the nanoparticles are

also required, however, to overcome a number of technical problems and possible

issues of toxicity.

16.5

Simulated Body Fluids

Artiﬁ cial materials implanted into bone defects might be encapsulated from time

to time by a ﬁ brous tissue, leading to their isolation from the surrounding bone.

Any improvement in bone - implant bonding to the tissues requires the correct

implant surface morphology and chemistry to generate a mechanical interlock and

good surface activity. Interactions between the bone and the implant will be con-

trolled with appropriate biological interactions.

542 16 Nanoceramics for Medical Applications

During the past three decades, many investigators have proposed that the essen-

tial requirement for an artiﬁ cial material to bond to living bone is the formation

of a bone - like apatite on its surface when implanted in the living body. In 1991,

Kokubo et al . proposed that in vivo apatite formation on the surfaces of many

biomedical materials could be reproduced in a SBF with ion concentrations almost

equal to those of human blood plasma. In essence, this means that the in vivo

bone bioactivity of a material can be predicted from the apatite formation on its

surface in SBF (Figure 16.7 ) [42] .

Hydroxyapatite layers can be easily produced on various organic and inorganic

substrates when submerged in SBF and indeed, in 1989, Kokubo and Takadama

[42] showed that, after immersion in SBF, a wide range of biomaterial surfaces

initiated very ﬁ ne crystallites of carbonate ion - containing apatite. Subsequently,

many reports have described the ability of osteoblasts to proliferate and differenti-

ate on this apatite layer. Based on these ﬁ ndings, other SBFs have been produced

in order to provide insight into the reactivity of the inorganic component of blood

plasma, and to predict the bioactivity of implants and bone scaffolds, as well as

other novel biomaterials. SBF has also been used to prepare bioactive composites

by forming HAp on various types of substrate.

The SBF solutions have been shown to induce apatitic calcium phosphate forma-

tion on any metal, ceramic, or polymer soaked in them. The SBF solutions, which

closely resemble Hank ’ s balanced salt solution ( HBSS ) [39] , are prepared with the

aim of simulating the ion concentrations present in human plasma. Hence, the

solutions are prepared with relatively low calcium and phosphate ion concentra-

tions (i.e., 2.5 and 1.0 m M , respectively), while the pH is adjusted to a physiological

value of 7.4 by using organic buffers (e.g., Tris or HEPES).

Typically, a SBF solution will have ionic concentrations of 142.0 m M Na

, 5.0 m M

, 1.5 m M Mg

, 2.5 m M Ca

, 147.8 m M Cl

−

, 4.2 m M HCO

3 −

, 1.0 m M HPO

2−

and 0.5 m M

2−

, with a pH of 7.4, all of these values being almost equal to those

of human blood plasma at 36.5 ° C. The SBF is usually prepared by dissolving

reagent - grade NaCl, NaHCO

, KCl, K

HPO

· 3H

O, MgCl

· 6H

O, CaCl

and

in distilled water and buffering at pH 7.4 with Tris(hydroxymethyl)ami-

nomethane ((CH

OH)

CNH

) and 1.0 M hydrochloric acid at 36.5 ° C.

Since their ionic composition is more or less similar to that of human blood

plasma, the HBSS or SBF formulations have only limited power with regards to

the precipitation of apatitic calcium phosphates. As a direct consequence, the

nucleation and precipitation of calcium phosphates from HBSS or SBF solutions

is rather slow. The time taken to achieve total surface coverage of a 10 × 1 0 × 1 mm

titanium or titanium alloy substrate immersed in a 1.5 × or 2 × SBF solution is

typically two to three weeks, with frequent (every 36 – 48 h) replenishment of the

solution [43] .

Among the metallic oxide gels prepared using a sol – gel method, those consist-

ing of SiO

, TiO

, ZrO

, and Ta

were found to have apatite formation on their

surfaces in SBF, as shown in Figure 16.7 . These results indicated that the Si – OH,

Ti – OH, Zr – OH, and Ta – OH groups on the surfaces of these gels were effective

in inducing apatite formation on their surfaces within the body environment

(Figure 16.8 ).

16.5 Simulated Body Fluids 543

Figure 16.7 Treatment of a titanium alloy surface with NaOH

to produce a bioactive surface on which hydroxyapatite

particles nucleate and grow quite readily within the SBF

solution [40] .

A variety of studies have been conducted using SBF solutions to deposit

apatite on both two - dimensional ( 2 - D ) and 3 - D scaffolds. For example, Wu and

coworkers [41] developed a novel bioactive, degradable and cytocompatible bredig-

ite (Ca

MgSi

) scaffold with a biomimetic apatite layer ( BTAp ) for bone - tissue

engineering. For this, porous bredigite scaffolds were ﬁ rst prepared using the

polymer sponge method. A BTAp was then applied to the scaffolds by soaking

them in SBF (pH 7.4) at 37 ° C for 10 days, with a solution volume - to - scaffold mass

ratio of 200 ml g

− 1

. After soaking, the scaffolds were dried at 120 ° C for one day,

such that bredigite scaffolds with BTAps were obtained. The porosity and in vitro

degradability of the BTAp scaffolds were investigated. Likewise, the osteoblast - like

cell morphology, proliferation and differentiation on BTAp scaffolds were evalu-

ated and compared with those of β - tricalcium phosphate ( β - TCP ) scaffolds. The

results showed that the bredigite scaffolds possessed a highly porous structure

with a large pore size (300 – 500 μ m). This biomimetic process mimics biominer-

alization and leads to the formation of a bone - like apatite layer on the scaffold

544 16 Nanoceramics for Medical Applications

surface. The obtained BTAp scaffolds also possessed a high porosity (90%) and

pore interconnectivity. When compared to β - TCP scaffolds, the cells on BTAp

scaffolds showed a higher proliferation rate and differentiation level.

Cromme and coworkers [44] investigated the activation of regenerated cellulose

2 - D model thin ﬁ lms and 3 - D fabric templates with calcium hydroxide, Ca(OH)

For this, the Langmuir – Blodgett ( LB ) ﬁ lm technique was applied to manufacture

model thin ﬁ lms using a trimethylsilyl derivative of cellulose (TMS - cellulose).

Regenerated cellulose ﬁ lms were obtained by treating the TMS - cellulose LB - ﬁ lms

with hydrochloric acid vapors. For the 3 - D templates, regenerated cellulose fabrics

were used, and the templates activated with a Ca(OH)

- suspension and subse-

quently exposed to 1.5 × SBF to induce the in situ formation of calcium phosphate

phases. The calcium phosphates were identiﬁ ed using Fourier transform infrared

( FTIR ) and Raman spectroscopy as highly carbonated apatite s ( CA ) lacking

hydroxyl ions. Such 3 - D fabric templates of regenerated cellulose covered with a

Figure 16.8 Apatite formation on silica, titania, zirconia, and

tantalum oxide gel - covered substrates within the SBF solution

[40] .

16.5 Simulated Body Fluids 545

biomimetic coating of apatite might be of particular interest for novel scaffold

architectures in bone repair and tissue engineering.

Lim and coworkers [45] noted that bone - like apatite could be more efﬁ ciently

coated onto the scaffold surface by using polymer/ceramic composite scaffolds

rather than polymer scaffolds, and by using an accelerated biomimetic process to

enhance the osteogenic potential of the scaffold. The creation of a bone - like,

apatite - coated polymer scaffold was achieved by incubating the scaffolds in SBF.

Apatite growth on the porous poly( D , L - lactic - co - glycolic acid)/nanohydroxyapatite

( PLGA/HAp ) composite scaffolds was signiﬁ cantly faster than on the porous

PLGA scaffolds. In addition, the distribution of coated apatite was more uniform

on the PLGA/HAp scaffolds than on the PLGA scaffolds. After a ﬁ ve - day incuba-

tion period, the mass of apatite coated onto the PLGA/HAp scaffolds incubated

in 5 × SBF was 2.3 - fold higher than on the PLGA/HAp scaffolds incubated in

1 × SBF. Furthermore, when the scaffolds were incubated in 5 × SBF for ﬁ ve

days, the mass of apatite coated onto the PLGA/HAp scaffolds was 4.5 - fold higher

than on the PLGA scaffolds. These results indicated that the SBF - initiated apatite

coating could be accelerated by using a polymer/ceramic composite scaffold and

concentrated SBF. It was reported that, when seeded with osteoblasts, the apatite -

coated PLGA/HAp scaffolds exhibited signiﬁ cantly higher cell growth, alkaline

phosphatase ( ALP ) activity, and mineralization in vitro compared to the PLGA

scaffolds coated only with HAp. In conclusion, the biomimetic apatite coating

could be accelerated by both introducing nucleation sites into polymer scaffolds

and using concentrated SBF. When seeded with osteoblasts, scaffolds with acceler-

ated apatite coating signiﬁ cantly enhanced cell growth, ALP activity, and miner-

alization in vitro .

The SBF method was used by Kolos et al . [46] to fabricate calcium phosphate

ﬁ bers for biomedical applications. A natural cotton substrate was ﬁ rst pretreated

with phosphorylation and a Ca(OH)

- saturated solution, and then soaked in SBF

of two different concentrations, namely 1.5 × and 5.0 × the ion concentration of

blood plasma. The cotton was then burned out by sintering the ceramic coating

at 950 ° C, 1050 ° C, 1150 ° C, and 1250 ° C, such that hollow calcium phosphate ﬁ bers

approximately 25 μ m in diameter and with a 1 μ m wall thickness were success-

fully manufactured. However, the 5.0 × SBF produced a thicker and more crystal-

line coat of greater uniformity. Kolos et al . further reported that osteoblastic cells

were able to cover the entire surface of the cotton ﬁ bers; more surprisingly, the

cell coverage seemed to be independent of the surface roughness and the ﬁ bers ’

Ca : P ratio.

Although of micron size rather than nanosize, a bioactive CHAp layer on cel-

lulose fabrics was developed by Hoffman et al . [47] . Nonwoven cellulose (regener-

ated, oxidized) fabrics were coated with CHAp using a procedure based on the

SBF method. For this, SBF with a high degree of supersaturation (5 × SBF) was

applied to accelerate the biomimetic formation of bone - like apatite on the cellulose

fabrics. After creating calcium phosphate nuclei on the cellulose ﬁ bers in an initial

5 × SBF with high Mg

and

HCO

−

concentrations, the cellulose fabrics were

additionally soaked in a second 5 × SBF which was optimized with regards to

546 16 Nanoceramics for Medical Applications

accelerated crystal growth by reduced Mg

and HCO

−

concentrations. The car-

bonated apatite layer thickness was increased from 6 μ m after a 4 h soaking in the

latter solution, to 20 μ m after 48 h. The amount of CO

2−

substituting

3−

in the

HAp lattice of the precipitates could be varied by changing the soaking time.

16.6

Nano - and Macrobioceramics for Drug Delivery and Radiotherapy

16.6.1

Nanobioceramics for Drug Delivery

For drug delivery, the primary aim is to target drugs to speciﬁ c sites within the

body, and to release them in a controllable fashion. However, for many current

delivery systems the guest molecules are often released upon dispersion of the

carrier/drug composites in water. This type of premature release is particularly

undesirable and problematic when the guest molecule (e.g., an anti - tumor drug)

is cytotoxic and might potentially harm healthy cells and tissues before being

delivered to the affected sites [48] .

In the case of ceramics, the critical pore and grain size may be varied from a

few nanometers up to microns in order to control the ease of delivery and disper-

sion of a material to the targeted area. A variety of nanoceramic drug - delivery

systems are currently undergoing clinical evaluation. In addition to reducing

toxicity to nondiseased cells, these systems have the potential to increase drug

efﬁ ciency, which translates to signiﬁ cant cost savings for the expensive drug treat-

ments that currently are being engineered. On the basis of their physical size,

nano drug - delivery systems also have the extraordinary characteristic of being

able to target and control drug release with very high precision.

Mesoporous silica nanoparticle ( MSN ) materials, such as mobile composition

of material s ( MCM ) - 41/48, can be synthesized by utilizing surfactants as structure -

directing templates to generate a range of mesoporous structures with high surface

areas ( > 900 m

− 1

), tunable pore sizes in the range of 2 to 20 nm, and uniform

pore morphologies. Recent breakthroughs in terms of morphology control and the

surface functionalization of MSN materials have resulted in a range of new materi-

als that can be used as stimuli - responsive, controlled - release delivery - carriers for

many biotechnological and biomedical applications. As with some other conven-

tional drug - delivery agents with high loading capacities (e.g., polymer and lipo-

somes), MSNs can encapsulate large quantities of drugs with various sizes, shapes,

and functionalities [46] . In contrast to many biodegradable polymeric delivery

systems, in which the loading of drug molecules requires organic solvents, the

molecules of interest can be encapsulated inside the porous framework of the

MSN by capping the openings of the mesoporous channels covalently with size -

deﬁ ned “ caps, ” which physically block the drugs from leaching out. Drug mole-

cules loaded into the pores may then be released by the introduction of “ uncapping

triggers, ” with the rate of release being controlled by the concentration of the

16.6 Nano- and Macrobioceramics for Drug Delivery and Radiotherapy 547

trigger molecules. Prior to uncapping, the capped MSN system exhibits a negligi-

ble release of drug molecules. This “ zero - release ” feature of a capped MSN delivery

system, along with an ability to tune the rate of release by varying stimulant con-

centrations, are important prerequisites for developing delivery systems with

many site - speciﬁ c applications, such as highly toxic anti - tumor drugs, hormones,

and neurotransmitters to certain cells types and tissues [48] .

An MCM - 41 - type MSN - based, controlled - release delivery system has been syn-

thesized and characterized using surface - derivatized cadmium sulﬁ de (CdS)

nanocrystals as chemically removable caps to encapsulate several drug molecules

and neurotransmitters inside the organically functionalized MSN mesoporous

framework. Lai and coworkers [49] studied the stimuli - responsive release proﬁ les

of vancomycin - and adenosine triphosphate ( ATP ) - loaded MSN delivery systems

by using disulﬁ de bond - reducing molecules, such as dithiothreitol ( DTT ) and

mercaptoethanol ( ME ), as release triggers. The biocompatibility and delivery efﬁ -

ciency of the MSN system with neuroglial cells (astrocytes) in vitro was demon-

strated. In contrast to many current delivery systems, the molecules of interest

were encapsulated inside the porous framework of the MSN by capping the open-

ings of the mesoporous channels with size - deﬁ ned CdS nanoparticles, so as to

physically prevent the drugs/neurotransmitters from leaching out.

Porous aluminosilicate ceramics were investigated by Byrne and Deasy [50] for

their potential to act as extended - release drug - delivery systems. The aluminosili-

cate pellets were obtained either commercially, produced by extrusion - spheroiza-

tion, or by cryopelletization. It was reported that each product had a highly

interconnected porous microstructure, with the porosity and pore - size distribution

being product - dependent. Drugs were loaded into the pellets using a vacuum

impregnation technique, with the concentration of the drug loading solution and

pellet porosity inﬂ uencing the loading obtained. Each product provided an extended

release of the incorporated drug, with the rate - determining step of release being

the diffusion of the drug from the porous pellet interior into the bulk dissolution

medium. Byrne and Deasy [50] showed that this rate was inﬂ uenced by the pellet

size, its porosity, pore - size distribution and porous microstructure, and by electro-

static interactions between the pellet surfaces and the drug. The solubility of the

drug in the dissolution medium and its molecular weight also inﬂ uenced the

release rate. It was concluded that porous aluminosilicate pellets represent a par-

ticularly versatile class of extended - release drug - delivery system, as the drug is

incorporated into the pellets after their production.

A new TiO

nanostructured bioceramic device was synthesized by L ó pez and

coworkers [51] , using a sol – gel process in order to control the pore - size distribution

and particle size. The objective was to obtain a constant drug release rate for anti -

epileptic drugs directly into the central nervous system ( CNS ). This method of

drug delivery, using small reservoirs, is very important in pharmaceutical applica-

tions, as it offers advantages such as the elimination of secondary effects, a long

duration of pharmacological activity, and protection of the drug against enzymatic

degradation or pH variations. Among the best - developed and most studied materi-

als, Titania has been shown to be an excellent candidate because of the possibility

548 16 Nanoceramics for Medical Applications

to manipulate both the structure and the number of OH groups. Point defects

were generated in the Titania network in order to obtain the desired interaction

between a highly polar drug and the Titania device. The device contained an anti-

convulsant drug, valproic acid ( VPA ), which could be released directly into the

temporal lobe of the brain, at a constant rate. During the initial stage of the syn-

thesis, VPA was added until a completely homogeneous solution was created; this

solution was then maintained under constant stirring until a gel was formed.

Under these conditions, the titanium dioxide underwent nucleation, whilst at the

same time it was restructuring around the VPA in such a way that the chemical

and polar properties of the anticonvulsant were preserved. The reaction was seen

to take place under well - controlled conditions of pressure and temperature, which

were purposely kept low. During the gelation period, both the electronic and

molecular properties of the material were preserved. Finally, a dry gel was obtained,

in which the anticonvulsant was occluded within the pore structure. The release

of VPA into the temporal lobe of the brain, and its effect on epileptic rats, was

observed by using the Kindling method. Several important conclusions were

drawn from these studies, notably that a Titania device charged with an anticon-

vulsant drug could be successfully implanted in the rat ’ s temporal lobe. Pore -

blocking with higher concentrations of VPA led to a fall in the initial rate of drug

release, while insertion of the device caused a drastic reduction in the animal ’ s

epileptic activity.

16.6.2

Microbioceramics for Drug Delivery

The particulate forms of ceramic materials have found application in both medical

and non - medical ﬁ elds. Particles in the form of microspheres are especially appli-

cable when treating tumors located in organs that are supplied by a single afferent

arterial blood supply. More traditionally, microspheres and nanostructurally modi-

ﬁ ed ceramics have been used in the targeted drug delivery of chemotherapeutic

and radiotherapeutic agents.

In recent years, the preparation of surface - modiﬁ ed hollow microspheres has

attracted considerable attention because of their unusual properties, notably their

large speciﬁ c surface area due to the nanolayer - modiﬁ ed surfaces, their low density

and their encapsulation properties. Consequently, these materials should be very

useful for novel applications such as drug - and protein - delivery systems. In certain

applications, the efﬁ cacy of microparticulate materials can be greatly improved if

they can act simultaneously as carriers for biologically active molecules. In this

sense, porous and surface - modiﬁ ed materials have an advantage, as they present

an additional surface area that greatly inﬂ uences the loading capacity and release

rates.

In dental clinical applications, such as the treatment of severe periodontitis,

where massive alveolar bone loss occurs, bone defect ﬁ lling and intensive systemic

long - term antibiotics administration are often required. Nanohydroxyapatite

microspheres intended for use as injectable bone ﬁ lling material or as enzyme -