HHS Public Access

Nanostructured Platforms for the Sustained and Local Delivery

of Antibiotics in the Treatment of Osteomyelitis

Vuk Uskoković

Advanced Materials and Nanobiotechnology Laboratory, Richard and Loan Hill Department of

Bioengineering, College of Medicine, University of Illinois at Chicago, 851 South Morgan St, #205

Chicago, Illinois, 60607-7052

Vuk Uskoković: [email protected]

Abstract

This article provides a critical view of the current state of the development of nanoparticulate and

other solid-state carriers for the local delivery of antibiotics in the treatment of osteomyelitis.

Mentioned are the downsides of traditional means for treating bone infection, which involve

systemic administration of antibiotics and surgical debridement, along with the rather imperfect

local delivery options currently available in the clinic. Envisaged are more sophisticated carriers

for the local and sustained delivery of antimicrobials, including bioresorbable polymeric,

collagenous, liquid crystalline, and bioglass- and nanotube-based carriers, as well as those

composed of calcium phosphate, the mineral component of bone and teeth. A special emphasis is

placed on composite multifunctional antibiotic carriers of a nanoparticulate nature and on their

ability to induce osteogenesis of hard tissues demineralized due to disease. An ideal carrier of this

type would prevent the long-term, repetitive, and systemic administration of antibiotics and either

minimize or completely eliminate the need for surgical debridement of necrotic tissue. Potential

problems faced by even hypothetically “perfect” antibiotic delivery vehicles are mentioned too,

including (i) intracellular bacterial colonies involved in recurrent, chronic osteomyelitis; (ii) the

need for mechanical and release properties to be adjusted to the area of surgical placement; (iii)

different environments in which in vitro and in vivo testings are carried out; (iv) unpredictable

synergies between drug delivery system components; and (v) experimental sensitivity issues

entailing the increasing subtlety of the design of nanoplatforms for the controlled delivery of

therapeutics.

Keywords

calcium phosphate; composites; controlled drug delivery; nanoparticle; osteomyelitis

I. INTRODUCTION

Osteomyelitis, the infectious inflammation of bone and one of the oldest documented

diseases, the earliest descriptions of which date back to Hippocrates (fifth century BC),

an illness particularly prevalent among the elderly, diabetics, children, and indigenes of

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Crit Rev Ther Drug Carrier Syst. 2015 ; 32(1): 1–59.

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Third World countries (Fig. 1a). Before the advent of antibiotics, the mortality rate because

of osteomyelitis was 25–45%. Although morbidity due to chronic bone infection has

drastically decreased from the pre-penicillin era, down to ~3% in the past 20 years,

it is still

high on the global scale, and treating the disease continues to be considerably challenging.

The incidence of osteomyelitis in the United States is 1–2%, but the disease is far more

prevalent in developing countries, as well as among particular patient populations:

approximately 1 in 5000 children, 1 in 1000 neonates, 1 in 250 patients with sickle cell

disease, 1 in 7 diabetics, and 1 in 3 patients with punctured foot.

4–7

Its comparatively low

prevalence can be explained by the fact that bone is an organ well protected from external

pathogens and is not readily prone to infection. The difficulty faced by invasive pathogens in

an attempt to colonize the bone is, however, directly proportional to the difficulty faced by

clinicians in ensuring the delivery of antibiotics to the site of infection and curing it. The

prevalence of chronic osteomyelitis among patients treated for at least one episode of acute

osteomyelitis is consequently high, in the range of 5–25%.

Strategies for improving the

therapeutic approach in the treatment of osteomyelitis have thus been explored for over a

century,

with a steadily increasing annual number of publications related to it—from 1 to

10 until 1944 to 100–300 from 1944 to 1974 to 713 in 2012 (US National Library of

Medicine), more than in any of the preceding years—going in step with the anticipated

increase in the number of cases of bone disease as the corollary of the aging population

worldwide (Fig. 1b). The number of hip and knee replacement procedures performed in the

United States has, for example, doubled in the past decade, whereas the number of the

reported cases of bone infection accompanying those procedures also has steadily increased

in proportion to the number of surgeries performed (Fig. 2). In spite of using aseptic

techniques and antibiotic prophylaxis, osteomyelitis is estimated to develop in 22–66% of

patients following orthopedic operations, and the corresponding mortality rate could be as

high as 2%.

This review describes (1) the pathologies that cause osteomyelitis; (2) the

traditional therapeutic approach to curing it; and (3) advanced therapeutic methods based on

the design of nanostructured platforms for the sustained and local delivery of antibiotics.

II. PATHOLOGIES AND THE DOWNSIDES OF THE TRADITIONAL CLINICAL

APPROACH

Osteomyelitis is mainly caused by pyogenic bacteria found in healthy oral flora, although

cases of infection caused by fungi are also common.

13–15

Bone infections caused by

Brucella suis,

Haemophilus influenzae,

Mycobacterium tuberculosis,

Mycobacterium

ulcerans,

and pox viruses,

20,21

as well as those whereby bone lesions are secondary to

Bacille Calmette-Guérin

or smallpox

vaccination, also have been reported in the

literature. Although many gram-negative and gram-positive bacteria were reported to have

caused osteomyelitis; the great majority of bone infections, however, is staphylococcal in

origin and mostly caused by a single bacterium: Staphylococcus aureus.

24,25

In addition to

S. aureus, S. epidermidis is another common cause of osteomyelitis; it is present in up to

90% of bone infections following intraoperative implantation of a foreign material. Because

most cases of osteomyelitis are caused by bacteria that reside on healthy skin and in healthy

oral flora, osteomyelitis is an illness often caused by a bizarrely small scratch or a bite where

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by body fluids become exposed to external pathogens, which then go on to induce septic

arthritis and/or osteomyelitis.

The onset of the infection induces an acute suppurative inflammation, and numerous factors

synergistically contribute to the necrosis of the hard tissues, demineralization of the bone,

and degradation of its collagen matrix: bacteria, pH change, local edema that accumulates

under pressure, vascular obstruction, and leukocyte collagenase.

As the infection

progresses locally, it extends to the adjacent osseous structures through the Haversian and

Volkmann canals, leading to an increased obstruction of vascular channels and necrosis of

more osteocytes in the lacunae. By the time the infection reaches the outer part of the cortex,

it has already caused an inversion of the periosteal blood flow and gained access to the

subperiosteal space, which results in a subperiosteal abscess and the formation of

involucrum, a layer of new bone grown from periosteum stripped from the original bone

under the pressure of pus. Figure 3 shows radiological images of cases of acute and chronic

osteomyelitis (the former came from a clinic and the latter from an animal model),

along

with involucrum formed around the area of necrotic infection and a periosteal reaction in the

proximal area of the bone, respectively.

Osteomyelitis is particularly prevalent in the facial skeleton because of its accessibility to a

variety of external pathogens and commensal microorganisms.

It also presents a major

complication following orthopedic and maxillofacial surgeries,

including even the most

routine dental extractions.

Although the resistance to infection of healthy bone is naturally

high, implants reduce it by a factor of 10

(i.e., the number of pathogens sufficient to cause

an infection is reduced from 10

to 10

). As a result, intraoperative introduction of bacteria

accounts for the largest number of osteomyelitis cases, with the hip being a particularly

common orthopedic site of infection. Timely treatment of osteomyelitis is required to

prevent its spread to new sites in the body and to avoid systemic osteonecrosis or unaesthetic

facial disfigurement in the case of maxillofacial infection. The typical treatment regimen for

bone infection consists of (1) intravenous administration of antibiotics lasting 2–6 weeks,

frequently followed by a 6-month course of oral antibiotics in the case of chronic infection;

and (2) surgical removal of bone that has undergone necrosis due to restriction of blood flow

by the formed abscesses.

32,33

Correspondingly, the major downsides of the conventional

therapeutic approach include (1) systemic administration of the therapeutic agent and its side

effects; (2) low concentration of the therapeutic agent around the site of infection,

potentially inducing resistance of the pathogen to the antibiotic therapy; and (3) irretrievable

bone loss that often requires the insertion of implants or prostheses as lasting bone

substitutes. Moreover, in the advanced stages of infection, when bone necrosis has become

significant, the blood supply to the infected area is inadequate and the lesion is largely

inaccessible to antimicrobial agents transported by the blood stream. All these downsides

provide strong incentives in favor of the development of appropriate carriers for the local

delivery of antibiotics in the treatment of osteomyelitis.

III. CLINICALLY AVAILABLE MATERIALS FOR LOCAL DELIVERY

Because of the apparent downsides of traditional therapy, including primarily the systemic

and repetitive administration of antibiotics whereby the therapeutic concentrations in the

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target area constantly fluctuate between toxic and ineffective, steps have been taken to

develop particulate carriers for the local and sustained delivery of antibiotics following their

implantation directly at the zone of infection. Ever since the pioneering research in this field

carried out in Europe in the 1970s,

34,35

poly(methyl methacrylate) (PMMA) beads, first

clinically applied in 1972, have been the gold standard for the local delivery of antibiotics to

bone cavities. Currently, there is no clinical alternative to PMMA as a local delivery carrier

for osteomyelitis since they are the only preloaded option approved by the US Food and

Drug Administration (FDA).

PMMA beads loaded with hydrophilic antibiotics, including gentamicin, ceftriaxone,

tobramycin, and vancomycin, have been used with experimental and clinical success in the

past.

37–41

Despite this, numerous limitations are associated with the use of PMMA beads.

First, they are not biodegradable and require a secondary surgical procedure for removal

after the antibiotic is released through their porous polymeric structure. Second, PMMA

beads and spacers exhibit burst release

that depletes the drug from the carrier and is

followed by a rather insubstantial release period that may be insufficient to maintain a

therapeutic concentration for the desired 3–4 weeks and may even promote antibiotic

resistance.

Although release kinetics could be extended by increasing the size of the beads

and increasing the polymerization time, burst release has seemed unavoidable so far.

Because the release is conditioned by the diffusion of the drug through the porous polymeric

network and microscopic cracks in the cement—and not by the degradation of the polymer

the elution profiles show broad variations depending on the nature of the antibiotic,

exhibiting intense burst release and a prompt decrease in concentration below the

therapeutic level in some cases.

45,46

Relatively low toxicity results from the absorption of methyl methacrylate monomers and

the associated carboxylesterase-mediated conversion of methyl methacrylate to methacrylic

acid,

47,48

whereas biofilm frequently forms on antibiotic-laden PMMA beads, hindering the

antimicrobial action.

49,50

Although products preloaded with gentamicin are available on the

market (Septopal), most clinically applied PMMA beads are loaded with the antibiotic just

before surgical insertion

(Fig. 4), which can lead to inconsistent release profiles.

This

nullifies the producer’s liability for the product and makes it therapeutically applicable only

with the patient’s consent.

Finally, a comprehensive clinical study has yet to prove that

PMMA beads are more effective than the systemic antibiotic delivery approach in treating

orthopedic infections.

No significant difference in the treatment success rate was typically

observed when debridement was followed by the implantation of antibiotic-containing

PMMA beads for local release or the prescription of systemic antibiotics.

The lower cost

of the therapy is often considered the only advantage of local delivery using PMMA

beads.

Consequently, a large population of clinicians is skeptical about the benefits of

local delivery in the management of osteomyelitis compared with the traditional approach

and resorts to the latter in their practice.

With no nonbiodegradable bone substitutes for load-bearing applications in sight for either

the biomedical device market or anywhere in clinical testing, for many decades now the

greatest potential among the bone engineers has been logically ascribed to bioresorbable

implants. However, the only currently clinically available bioresorbable alternative to

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PMMA beads, predominantly used outside the United States, is calcium sulfate cements.

Unlike PMMA, whose monomeric absorption has previously caused intraoperative

cardiopulmonary complications during arthroplasty,

calcium sulfates are nontoxic and

inexpensive and have been successfully used as a drug carrier in the treatment of

osteomyelitis.

59,60

In addition to being used since the late 19th century as a bone filler in

their hemihydrate form, also known as plaster of Paris,

calcium sulfates have been applied

in reparative dentistry for maxillary sinus floor augmentation

and for the repair of

periodontal defects

and root perforations.

In addition to their exceptional softness and

poor handling features, the main downside is that they are resorbed rapidly, in a matter of

weeks—faster than the bone ingrowth rate—which can lead to mechanical implant failure.

An ideal bioresorbable implant provides a mechanical support that is gradually transferred to

the newly formed bone, a requirement that calcium sulfates do not satisfy. They can also

cause severe drainage at the wound site after the surgical implantation,

as well as the

formation of a fibrous gap in the area where the slowly ingrowing bone replaces the rapidly

resorbing cement,

the same effect that is expected to result from the use of aragonite

calcium phosphate phases, such as tricalcium

or dicalcium

70,71

phosphates, as bone fillers;

these are more soluble than hydroxyapatite, the calcium phosphate constituent of bone. Also,

as a result of their relatively fast degradation in the body, the concentration of the antibiotic

at the target site and its mean blood serum concentration over the first month following

implantation are lower when compared with hydroxyapatite.

For this reason, and in view

of the fact that calcium sulfates have not led to therapeutic outcomes any better than those of

PMMA implants,

their use as an ideal bioresorbable delivery vehicle for antibiotics and a

void filler in bony defects has been questioned.

IV. ADVANCED DRUG DELIVERY PLATFORMS IN THE RESEARCH STAGE

As noted earlier, two main disadvantages of the traditional treatment of osteomyelitis

include (1) systemic distribution of the therapeutic agent and (2) the need for surgical

removal of necrotic bone. Options for sustained antibiotic release that can ensure high local

concentrations and low serum concentrations of the drug

already exist, and progress in

terms of promoting the osteogenic activity of the carrier is expected in the future. Whereas

local and sustained release of the drug could overcome the need for prolonged oral and/or

intravenous antibiotic therapies, the induction of osteogenesis by the carrier itself or the

growth factors released from it could eliminate or at least minimize the surgical removal of

affected bone, along with the frequent skeletal deformations and unaesthetic physical

disfigurements it entails. Patients with diabetic neuropathy are prone to developing

osteomyelitis of the forefoot, which often leads to minor amputation

; with the

development of osteogenic carriers that could revitalize the diseased bone, however, such

clinical cases could be coped with in a manner less traumatic for the patient. For example, in

parallel with the drug release process, the particles may decompose, dissipating their

osteogenic contents and thus fostering the bone healing process and natural restoration of the

portion of bone damaged by the pathogen. The therapeutic approach to treating

osteomyelitis would clearly yield a whole new dimension by using one such osteogenic drug

delivery platform. After all, with an ideal therapeutic agent serving a dual purpose of (1)

eliminating the source of illness and (2) revitalizing the organism, the conception of drug

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delivery carriers that exhibit simultaneous bactericidal and osteogenic performance is

natural.

A. Calcium Phosphates

Calcium phosphates occupy a special place among the biodegradable drug carriers of

antibiotics in bone repair. They have been traditionally considered a convenient choice for

the synthetic substitute of hard tissues because of their excellent biocompatibility,

osteoconductivity, lack of cytotoxicity, nonimmunogenicity, and sufficient loading

capacities, thanks to which hydroxyapatite, their least soluble phase, has been used as a

chromatographic adsorbent of proteins,

77–79

nucleic acids,

80–82

and microorganisms.

Excellent adsorption properties of hydroxyapatite are the result of its positively charged

surface Ca

ions engaging in an anion-exchange interaction with deprotonated carboxyl

groups of proteins and the negatively charged PO

3−

groups engaging in a cation-exchange

interaction with protonated amino groups of proteins.

Moreover, hydroxyapatite possesses

different net charges on the a and c planes of its hexagonal crystal lattice—positive and

negative, respectively,

which renders it effective in the crystallographically selective

binding of multiple molecular entities. Other variations of hydroxyapatite, such as

carbonated apatite

and biphasic calcium phosphate,

possessed an even greater protein

adsorption capacity, given an identical particle size and specific surface area, which was

hypothesized to be due to their greater solubility, which increases the ionic strength in the

medium and the surface exposition of the polar residues of proteins, thus increasing the

binding efficacy.

Hydroxyapatite also has been used as an amphiphilic stabilizer in

Pickering emulsions, suggesting its ability to interact with both hydrophilic and hydrophobic

compounds.

Unlike PMMA, calcium phosphates are fully bioresorbable, and the rate of their degradation

could be tentatively tuned by controlling the phase composition of the compound.

Namely,

as can be seen in Table 1, calcium phosphates can adopt a variety of stoichiometries,

covering a range of solubility product values, from 0.07 for anhydrous and monohydrous

monocalcium phosphates to 10

−7

for monetite and brushite to 10

−25

for α-tricalcium

phosphate to 10

−117

for hydroxyapatite.

Of course, because more ionic species exist in the

stoichiometric formulas of the less soluble phases, the difference in solubility is of a lesser

magnitude than that in the solubility product, amounting to approximately 6 · 10

, 1.6 · 10

and 8.3 times higher solubility for monocalcium phosphate, monetite, and α-tricalcium

phosphate, respectively, compared with hydroxyapatite (0.3 mg/dm

) in water at 37°C and

at a physiological pH.

Particle size presents an important consideration in the design of most optimal degradation

and release profiles, and nanosized calcium phosphates have proved to be far more

advantageous than the microsized ones,

a natural consequence of the fact that bone itself

contains apatite particles with nanosized dimensions

(20 × 10 × 2 nm, on average

Furthermore, porosity that is controllable via sintering at elevated temperatures could be

used to vary the degradation rate in vivo within a wide window of values; again, nanosized

and fully dispersed hydroxyapatite is highly bioresorbable and the denser formulations are

resorbable to a significantly lesser degree,

leading to hypotheses that nonporous, sintered

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hydroxyapatite blocks should be stable in biological milieus for centuries.

In this case the

drug release rate tends to be directly proportional to the resorption rate, that is, significantly

higher for the more porous calcium phosphate microstructures.

Stoichiometry of single-

phase compositions, implant geometry, ionic substitutions, crystallinity, and macro- and

microporosity are other factors known to greatly affect the degradation rate of calcium

phosphates in vivo.

98–100

When self-setting calcium phosphates pastes are used, the powder-

to-liquid ratio, initial viscosity, pH, and the presence of additives, such as crystallization

seeds, inhibitors, or dispersants, are additional factors that influence the hardening

properties, the degradation kinetics, and the rates of resorption and new bone ingrowth,

101

which usually range anywhere between 3 months and 3 years.

102

Calcium phosphates are also a component of the mineral phase of hard tissues, which makes

them a natural candidate for bone-filling drug carriers. With bone acting as a natural

reservoir for calcium and phosphate ions,

109

any excessive amounts thereof could be

regulated in favor of new bone growth. Calcium and phosphate ions released upon the

degradation of these compounds can also stimulate osteoblastic differentiation

110,111

and

proliferation

112

and be used as ionic ingredients for the formation of new bone. Another

advantage of calcium phosphates is that they could be sterilized by a variety of techniques,

including γ-irradiation, gas plasma, supercritical carbon dioxide, or even steam autoclaving

(in the case of hydroxyapatite), without causing adverse effects to their structure and

properties. By contrast, in general there is currently no established sterilization procedure for

polymers that does not modify their structure to some degree, due to (1) physical

deformations and chemical changes—scission and cross-linking that occur upon

autoclaving,

113

alongside practically inevitable degradation of an encapsulated drug

114

; (2)

surface chemistry modifications that occur upon the application of ethylene oxide, hydrogen

peroxide, or ozone

115

; (3) bulk structural changes and a decrease in the molecular weight

that occur during γ-irradiation,

116

while a difficult regulatory path is posed before novel or

nontraditional sterilization methods.

Calcium phosphates are also relatively easy to prepare in a variety of morphological

forms,

117

although not at a particle size below 20 nm, as is the case with metals. Different

calcium phosphate particle morphologies possess different bioactivities,

118,119

which allows

for the optimization of their biological response by means of controlling morphological and

specific crystal face exposition. Calcium phosphates are also naturally precipitated in a

nanosized form, and the use of nanoparticulate calcium phosphates could be considered as a

win–win solution in the quest for simultaneous bactericidal and osteogenic properties.

Namely, the drug adsorption efficiency is directly proportional to the specific surface area of

the adsorbent and inversely proportional to the particle size.

120

The large surface area of

nanosized calcium phosphates thus increases their drug-loading capacity and makes them a

more effective bactericidal agent.

121

At the same time, nanosized calcium phosphates

possess higher bioactivity than their microsized counterparts,

122,123

an insight that is natural

in view of the nanosized dimensions (30 × 20 × 2 nm)

124

of mineral particles in bone. Last

but not least, calcium phosphates are one of the safest nanomaterials evaluated for toxicity

so far.

125

Figure 5a displays round hydroxyapatite nanoparticles obtained by precipitation

from alkaline aqueous solutions and highlights their ability to capture large amounts of drug

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molecules in the pores between the particles upon desiccation at low pressure.

Mechanistically similar intraporous loading of hydroxyapatite with a drug was reported

earlier for isepamacin sulfate, an aminoglycoside antibiotic.

126

The same effect of extended

release could be achieved by compacting the antibiotic-loaded calcium phosphate powders

under pressure.

127

PerOssal, a commercial mixture comprising 51.5% nanocrystalline

hydroxyapatite and 48.5% calcium sulfate, for example, relies on such compaction of

nanoparticles to ensure sustained release of antibiotics.

128

1. Concerns Pertaining to the Use of Calcium Phosphates—The application of

calcium phosphate particles as drug delivery carriers naturally has its downsides, and the

main one comes from their difficult surface functionalization. This is, in part, the effect of

their ionic nature, which dictates that the surface layers undergo rapid reorganization via

dissolution/reprecipitation phenomena in ionic media. As evidence of this effect, ζ potential

of hydroxyapatite particles changed with the immersion time, indicating an exchange of ions

across the interface layer and its restructuring following local changes in the solvent

medium.

129,130

Despite the presence of calcium, phosphate, and hydroxyl ionic groups on

the particle surface, which, in theory, would allow the binding of an array of functional

groups, the intense ionic exchange between the particle surface and its ionic milieu renders

this approach inoperative for dispersed particles. Compared with calcium phosphate

nanoparticles, silanol groups on the surface of silica nanoparticles offer greater stability and

more facile functionalization with organic molecules, having the same role as monolayers of

thiol groups chemisorbed on the surface of silver, copper, or gold

132

and carboxylic or

phosphonic acid moieties on the surface of metal oxides or quantum dots. Their downside,

however, is an uncertain fate in the body and an array of inflammatory and oxidative

stresses possibly induced in it, ranging from mitochondrial dysfunction to genotoxicity to

pulmonary congestion to hepatocyte necrosis.

133–135

Unlike polymeric materials (e.g., hyaluronic acid), whose viscosity could be controlled to a

greater degree by means of chemical or photochemical cross-linking, thixotropic calcium

phosphate cements exhibit a far narrower window of setting rates, which significantly limits

the flexibility of their surgical handling. Variations in the concentration of plasticizing

additives, liquid-to-solid ratio, particle size and sphericity, and ionic strength of the liquid

phase have all been studied in a search for the optimal conditions for the fabrication of

injectable but cohesive calcium phosphate pastes and putties.

136

On the other side of the

spectrum, occupied by solid and strictly implantable materials, nonsintered calcium

phosphates in particular—which are strong but fragile and most interesting for drug delivery

applications are hardly formable and also are difficult to surgically attach to bone with

screws and intramedullary rods, for which reason they are often combined with a tougher,

more ductile organic phase to mimic the mechanical properties of bone itself.

137

Also, as a

consequence of uncontrolled ripening in the nucleation and crystal growth stages,

monodisperse calcium phosphates are difficult to prepare in a broad array of sizes, which

poses obstacles to systematic studies of the effect of calcium phosphate particle size on

bioactivity.

138

Last but not least, the most important disadvantage of calcium phosphates is that they have

little or no ability to be loaded with organic molecules via intercalation, which limits the

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loading mechanism to physisorption only and makes prolonged release difficult to achieve.

Namely, burst release typically results when the drug is adsorbed on the surface of the

carrier, and while, on one hand, this effect is favorable in terms of ensuring that the minimal

inhibitory concentration for the given pathogen is exceeded, it can also deplete the carrier

from the antibiotic and make its further release therapeutically ineffective. Still, a

tremendous difference between microsized and nanosized calcium phosphate particles was

found: Whereas the concentration of vancomycin released from the former was below the

detection limit 10 days after the implantation, nanoparticles of the same composition were

able to sustain the therapeutic level of release for up to 6 weeks.

139

Extended release from

porous calcium phosphate cements and its therapeutic effects in vivo were confirmed on

numerous other occasions.

140–143

Finally, because of relatively low ζ potentials (<15 mV on

the absolute scale), calcium phosphates form sols of low stability; simple and rapid

precipitation procedures for their formation in the low crystalline and nanoparticulate form,

on the other hand, enable them to be prepared before their clinical application.

2. Calcium Phosphate as an Intrinsically Osteoinductive Material—Calcium

phosphates have been generally considered as osteoconductive materials in the sense that

they support bone growth on them, although their ability to upregulate the expression of

osteogenic markers and boost osteoblastic differentiation, making them osteoinductive, too,

has been reported on numerous occasions.

144–146

The addition of growth factors, such as

bone morphogenetic proteins (BMPs), also has made calcium phosphates

osteoinductive,

147,148

although the same osteoinductive effect achieved by BMP-2 on

human mesenchymal stem cells was accomplished by nanosized hydroxyapatite particles.

149

In a corresponding study composites for the delivery of recombinant human BMP-2 (rH-

BMP-2) to mice and rabbits, comprising poly(D, L-lactic acid), p-dioxanone, polyethylene

glycol (PEG), and β-tricalcium phosphate needed less of the BMP than the same composites

that excluded hydroxyapatite from their composition to induce the same osteopromoting

effect and new bone formation.

150,151

Another study demonstrated that the expression of

BMP-2 in human periodontal ligament cells increased upon stimulation with nanosized

hydroxyapatite.

152

Optimization of substrate topography was able to yield the same

differentiation–induction effect as the chemical differentiation agents in the transformation

of mesenchymal stem cells to osteoblastic ones,

153

and a similar approach that could be

applied to ensure induced osteogenic response of bone cells without the use of expensive

growth factors would be great news, especially since bone infection is an illness known to

be particularly prevalent among patients in the Third World countries, for whom

affordability presents a vital feature of a marketed drug. This does not even consider that the

use of rHBMP-2 in bone augmentation procedures has induced ectopic bone formation,

osteolysis, pseudoarthrosis, inflammatory reactions in soft tissues, increased risk of

malignancies, and other adverse effects,

154,155

raising significant concerns over its safety in

the recent years.

156

The combination of rHBMP-2 with calcium phosphates has, however,

mitigated these adversities associated with the direct infusion of the given growth factor or

its delivery using organic carriers.

157

The naturally bactericidal citrate ion, accounting for

5.5 wt% of the organic content of bone, where it coats hydroxyapatite crystals at 0.5

molecules/nm

and stabilizes them in the collagen matrix,

158

increases in concentration in

parallel with the differentiation of mesenchymal stem cells into osteoblasts

159

and has been

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proposed as an alternative to BMPs in view of the ability of composites comprising

hydroxyapatite, in combination with polymers based on citric acid, to facilely regenerate

necrotic bone.

160

To render calcium phosphates as a base for an authentically osteogenic material by including

cell components capable of bone production, such as osteoprogenitor cells or differentiated

osteoblasts, however, the formation of porous scaffolds based on calcium phosphates is

needed, comprising a difficult but not impossible task.

161–164

For example, with the addition

of only 3 vol% gelatin, electrospinning, the method traditionally used to obtain polymeric

scaffolds, could be used to prepare calcium phosphate scaffolds as well.

165

Biomimetic

methods based on the usage of porous biological hard tissues as casting molds for the

synthesis of structurally similar inorganic scaffolds also have recently gained popularity.

166

In addition, a combination of self-setting pastes and porogens, such as mannitol

crystals,

167,168

pectin,

169

hydrogen peroxide,

170

inorganic crystals,

171

surfactants,

172,173

poly-(D,L-lactide-co-glycolide) (PLGA),

174,175

oils,

176

or other hydrophobic compounds,

was also used to create macroporous calcium phosphate formulations. Calcium phosphate

nanoparticles were successfully incorporated in polymeric,

177

collagen,

178

or carbon

nanotube

179

scaffolds with the purpose of promoting greater adsorption of adhesive serum

proteins and inducing bone growth. Simple admixing of microsized polymeric spheres into

calcium phosphate cements is another method used to produce macroporosity sufficient to

provide a proliferation milieu for host cells after the degradation of the polymeric phase.

180

3. Prospect of Ion-Substituted Calcium Phosphates—By affecting their lattice

parameters, crystallinity, and the solubility product, ionic substitutions in calcium

phosphates seem to have a large effect on a range of their physicochemical

properties.

181–183

While geological apatite can accommodate half of all the elements of the

periodic table in its crystal lattice,

184

biological apatite contains about a dozen different ions

as impurities, which has provided a rationale for the expected improvement in the biological

response to ion-substituted calcium phosphates.

185

Substitution of Ca

with K

, Na

, or

other alkali ions can, for example, increase the solubility of hydroxyapatite beyond that of

tricalcium phosphate.

186

Like Na

, Mg

is an ion that inhibits the nucleation of

apatite.

187,188

However, it is also the ion for which bone is the biggest reservoir in the body

and whose deficiency logically reduces bone growth,

189

explaining numerous attempts to

augment existing calcium phosphate formulations by doping them with Mg

190,191

Together with Mg

, Zn

has been found in subnormal concentrations in osteoporotic

patients, suggesting the vital role of these two cations in proper bone remodeling.

192,193

Because of the essential role of Zn

in the production of more than one bone growth

protein, including the zinc finger containing transcription factor Osterix,

194

the deficiency of

this micronutrient was proven to have detrimental repercussions on bone development, as

well,

195

which is another argument in favor of its incorporation into calcium phosphates

designed for bone substitutes. Zinc-substituted hydroxyapatite containing 1.6 wt% of Zn

possessed a more pronounced antibacterial effect against S. aureus compared with pure

hydroxyapatite.

196

Selenium is another element with strong antimicrobial properties that has

been introduced to carbonated hydroxyapatite via CO

2−

→ SeO

2−

substitution, with the

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resulting material being able to inhibit the formation of Pseudomonas aeruginosa and S.

aureus biofilm on its surface.

197

Silicon (Si) and strontium (Sr) are present in newly formed bone in the amounts of 0.5

198

and 0.03 wt%, respectively, and a more viable biological response was detected upon the

implantation of Si-doped and Sr-doped hydroxyapatite compared with pure

hydroxyapatite.

199–201

The most probable reason for this lies in the osteopromotive

properties of Si and Sr ions per se; Si has been demonstrated to increase bone mass density

and angiogenesis during new bone growth,

202

whereas Sr upregulates the expression of the

osteoblastic protein osteoprotegerin, which inhibits the production of RANKL and hinders

the differentiation and activation of osteoclasts.

203

Incorporation of either of these two ions

in the crystal lattice of hydroxyapatite increased the degradation of the compound in

vitro.

204,205

Vanadium is another element critical for healthy bone development because of

its ability to stimulate mineralization of collagen and proliferation of osteoblasts,

206

but the

bioactivity of vanadium-doped calcium phosphates

207

has yet to be assessed.

Calcium phosphates are able to sequester heavy ions from the environment, such as Pb

and As

, which is why they have been used as adsorbents in water purification.

208

Calcium

phosphate particles could thus be easily doped with Eu

, Tb

, Gd

, La

, or other

lanthanides and be made luminescent and used for imaging applications.

209,210

Hydroxyapatite labeled with

Tm,

125

Yt,

153

Sm, or

H radionuclides could also be

considered for simultaneous bone substitution and imaging applications.

211–213

Doping

hydroxyapatite with alkaline earth metals and magnetic elements, such as cobalt

214

iron,

215–217

yielded other impure forms of calcium phosphate that have been intensively

researched for their unique bioactive properties.

218–220

Superparamagnetic hydroxyapatite

obtained by doping with approximately 10 wt% Fe

/Fe

was hailed as a far less toxic

alternative to magnetite when used as a heating material for hyperthermia-based bone cancer

therapies.

221

Finally, carbonated hydroxyapatite, structurally similar to its biomineralized

form, has been frequently demonstrated to be superior in terms of its bioactivity compared

with its stoichiometric, noncarbonated counterpart.

222,223

Explored alternatives to calcium phosphates and the two aforementioned materials in actual

clinical use, PMMA and calcium sulfate, include mainly various polymeric materials,

bioactive glasses, liquid crystals, collagen, and titanium nanotubes; these are discussed in

the sections that follow.

B. Synthetic Biodegradable Polymers

Synthetic biodegradable polymers proposed as potential antibiotic carriers in the site-

specific treatment of osteomyelitis are predominantly poly(α-hydroxy esters).

224–226

Among

them, poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), PLGA,

227

and poly(ε-

caprolactone)

228

(PCL) have been studied most. All of these compositions have a proven

history of encapsulating arrays of both hydrophilic and hydrophobic compounds, including

antibiotics,

229

and enabling their sustained, first-order release over prolonged periods of

time.

230

While being formable in situ and capable of fitting practically any shape of a bone

defect to be filled, they also allow for fine-tuning of their mechanical and degradation

properties via control over their chemical structure, including parameters such as molecular

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weight, crystallinity, cross-linking ratio, and end-group identity. For example, the

decomposition kinetics of PLGA could be easily controlled by varying the lactide-to-

glycolide ratio and made to match the rate of new bone formation; namely, while PLLA has

a relatively lengthy degradation time scale, ranging from 3 months to over a year, depending

on the molecular weight, crystallinity, and other physicochemical factors, a gradual increase

in PGA content shortens this degradation to a matter of weeks for PLGA 50:50 as a result of

the decreased crystallinity and higher hydrolysis rate of PGA, after which the crystallinity

and resistance to degradation increase again at higher PGA contents, producing the

characteristic U-shaped curve (Fig. 6a).

231,232

An example of how sensitive the kinetics of

degradation and drug release could be to cross-linking ratio is shown in Fig. 6b; whereas 0.5

% of cross-linking in an acrylic hydrogel completes release in less than 5 hours, 1% of

cross-linking promotes sustained release over a period of 8 days.

233

Other biodegradable

synthetic polymers developed and tested as potential carriers of antibiotics in the treatment

of osteomyelitis include poly(trimethylene carbonate)

234,235

; polyamide fibers

236

;

polyhydroxyalkanoates, e.g., poly(3-hydroxybutyrate-co-3-hydroxyvalerate)

237

; and

polyanhydrides, e.g., poly(sebacic anhydride)

238

; poly(sebacic-co-ricinoleic-ester-

anhydride)

239

; or Septacin,

240

a copolymer of dimeric erucic acid and sebacic acid.

Polymeric composites are also the subjects of intense research. For example, a layer-by-

layer technique was used to grow multilayered polyelectrolyte films incorporating

gentamicin and comprising a cationic poly(β-amino ester) and anionic poly(acrylic acid) on

top of nondegradable poly(ethyleneimine) and poly(sodium 4-styrenesulfonate). Despite the

fact that more than two-thirds of the drug content were released in the first 3 days, the

implants were successful in treating S. aureus infection in a rabbit bone model.

241

1. Concerns Pertaining to the Use of Aliphatic Polyesters—Although poly(α-

hydroxy esters) have been successfully used in bone tissue engineering since the early

1990s,

242–244

there exists a concern that their acidic degradation products may favor

bacterial growth and promote hard-tissue resorption and bone mass loss,

245,246

effects

experimentally evidenced in the past. In spite of the supposed safe inclusion of the

byproducts of the degradation of PLLA-based polymers in the metabolic cycles of the host

organism (e.g., lactic acid is secreted by osteoclasts to resorb bone and is also one of the

compounds in the Krebs cycle), chronic inflammation has often resulted as a response to

their implantation in bone tissue engineering.

247–249

Another concern is that this

acidification effect may render rather ineffective antibiotics whose antimicrobial

effectiveness exists within only a narrow window of pH values. A decrease in pH from 7.4

to 5.5, for example, has led to a 16-fold increase in the minimum inhibitory concentration of

clindamycin with respect to S. aureus.

250

Poly(α-hydroxy esters) also lack the mechanical

properties required for load-bearing applications; PLGA, combining the adsorptive stability

of PLA with the mechanical strength of PGA, is the most favored and thus the most

researched option with respect to this intrinsic drawback.

C. Gels and Bioderived Polymers

Aqueous monoolein gels are an example of a liquid crystal system that was used to deliver

gentamicin sulfate for 3 weeks without the burst effect.

252

A mannosylated poly-

phosphoester gel with the capability of targeting macrophages and releasing the antibiotic

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payload in a site-activated manner, that is, only after being degraded by the bacterial

enzymes, was developed.

253

Use of the osteointegrating effects of calcium phosphates was

attempted by incorporating them into cubic liquid crystals of gentamicin–mono-olein–water

formulations.

254

Various combinations of calcium phosphates with gelatins, that is, mixtures

of peptides and proteins resulting from partial degradation of collagen, also were

investigated.

255,256

Among other bioderived polymers, some have been used to encapsulate

antibiotics, such as albumin

257,258

or dextran,

259

but have not been reported in bone-related

experimental trials, except in combination with more mechanically stable, inorganic phases.

Albumin coatings around allografts, for example, improved cell adhesion and

proliferation

260

and enhanced bone healing,

261

whereas the use of dextran as a porogen in

PMMA beads boosted the release of vancomycin, daptomycin, and amika-cin.

262

Conversely, silk fibroin coating around PCL microspheres managed to reduce the initial

burst release of vancomycin and extend the timeframe of its release.

263

Silk–alginate

copolymers are particularly interesting because of their tunable stiffness as the function of

the silk-to-alginate ratio and the concentration of the crosslinker,

264

but they have not been

used yet for the controlled delivery of antibiotics. Other natural polysaccharides, such as

chitosan,

265–267

pectin,

268

amylose,

269

alginate,

270

and hyaluronic acid,

271

have been both

used for the controlled release of antibiotics in vitro and tested as a component of

antimicrobial bone grafts in vivo. Pectin microspheres, alone and in combination with

chitosan, were used to encapsulate ciprofloxacin and were more effective in treating

osteomyelitis than intramuscularly administered antibiotic in a rat model.

272

A cross-linked

amylose starch matrix loaded with ciprofloxacin prevented and eradicated infection more

effectively than oral ciprofloxacin treatments in dogs with an infected femur.

273

Vancomycin encapsulated within alginate beads and distributed in a fibrin gel scaffold was

used to treat infected tibiae in rabbits.

274

Still, the most researched among bioderived

polymers as a potential carrier of antibiotics in the treatment and prevention of orthopedic

infection is collagen.

D. Collagen Sponges

Outside the United States in the 1980s, collagen sponges, also known as fleeces, began to be

used as the major alternative to PMMA beads for the local delivery of antibiotics. Their

application has been justified by a moderate number of clinical and in vivo studies.

275

For

example, compared with PMMA beads, sponge-like collagen carriers of gentamicin were 7

times more effective in reducing the bacterial colony count in the treatment of osteomyelitis

caused by S. aureus in the tibiae of rats.

276

Also, the placement of gentamicin-eluting

collagen fleece around the fixation plate during the surgical treatment of open bone fracture

prevented surgical site infection from occurring and promoted bone union in a large

population of patients.

277

In fact, rather than as a bone filler, collagen has been mostly used

as a material for postsurgical prophylaxis in the treatment of infectious disease.

1. Concerns Pertaining to the Use of Collagen—In spite of (1) the viable tensile

strength of collagen, (2) its ability to foster cellular attachment, and (3) the fact that collagen

sponges have been successfully used in the past,

278

the choice of antibiotics in their

clinically applicable versions has been limited to gentamicin only, alongside other

disadvantages that collagen intrinsically possesses. The main problem associated with the

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application of collagen and its derivatives as bone fillers comes from the intrinsic

immunogenicity of the collagen molecule,

279

namely, most of it is xenogenic in origin

because it is difficult to obtain directly from a patient, and recombinant technologies and the

methods to extract the immunogenic, telopeptide portion of collagen molecules are not only

of limited availability but also lead to reduced bioactivity of the protein.

280

Although

collagen has been successfully applied topically, for example, in biodegradable sutures and

as a prophylactic wound dressing carrier of antibiotics,

281–283

its mere subcutaneous

epithelialization may lead to undesired immunogenic or antigenic responses.

284,285

While it

can lead to antigenic and inflammatory responses, collagen is also typified by comparatively

uncontrolled degradation and drug release rates in the body.

286,287

For this reason, a

combination of collagen sponges with other polymers has been used to render more

sustained drug release profiles. One such composite material enriched with chitosan

microspheres and delivering recombinant human BMP-2 considerably outperformed a pure

collagen sponge loaded with the same growth factor in terms of new bone growth

enhancement, bone/implant integration, and the duration of drug release.

288

E. Silicon-Based Materials

Porous bioactive glass scaffolds loaded with ceftriaxone demonstrated a higher local

concentration of the antibiotic 6 weeks after the implantation compared with a parenteral

treatment composed of two injections per day.

289

Silicate-to-borate replacement in bioactive

glasses produced materials that also were used for the controlled delivery of vancomycin or

teicoplanin and repair of infected bone in rabbits.

290,291

Partial substitution of PO

3−

groups

of hydroxyapatite with SiO

4−

species resulted in a calcium phosphate–based glass ceramic

able to release vancomycin in a sustained manner over 2 weeks after cross-linking with

chitosan.

292

The addition of Ag

ions to phosphate-based glasses led to their sustained

release and bactericidal effect against S. aureus biofilms.

293

Further research will, however,

be necessary to show whether such antibiotic-free methods are capable of acting against

severe bone infections. As far as silica-containing materials are concerned, xerogels

obtainable from a solgel process were used to encapsulate and ensure the prolonged release

of vancomycin, with the water-to-alkoxysilane molar ratio being discerned as a parameter

for the control of release kinetics.

294

Zeolites, microporous aluminosilicates with

pronounced (1) antibacterial,

295

(2) adsorptive,

296

and (3) bone-protective dietary

297

properties, inhibited osteoclast-mediated bone resorption in vitro,

298

but their application as

a component of bone fillers in combination with calcium phosphates or other

osteoconductive phases is still a largely unexplored area. Metal-organic frameworks,

mesoporous materials structurally related to zeolites,

299

have been proposed as potentially

efficient drug delivery carriers,

300

including in applications that pertain to bone

regeneration,

301

which, however, they have yet to be tested for. Their main weakness is

rapid degradability in aqueous media, and structural variants with increased stability in

water are being intensively sought.

302

F. Metals

The widespread rise in the resistance of common pathogens to organic antibiotics has led to

a greater degree of consideration of the use of metals to prevent or treat infection,

303

with

many of them, such as silver-impregnated fabrics used as prophylactic dressings during

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wound healing,

304

regularly applied in the clinic. Because of its antimicrobial properties,

305

the first metal proposed for use in the treatment of osteomyelitis was silver; be it alone or in

the form of nylon wire composites, clinical testing resulted in a 65% success rate and no

evidence of postoperative argyria.

306

Titania nanotubes formed by electrochemical

anodization on the surface of titanium nanowires used for bone fixation were capable of

being loaded with gentamicin and releasing it over a period of 2 weeks.

307

In another study,

however, the same combination of gentamicin and TiO

nanotubes with an 80-nm diameter

and 400-nm length led to prompt release of the drug in only 1–2 hours, but it still reduced

the adhesion of Staphylococcus epidermis on the surface compared with pure titanium.

308

Surface etching and anodization parameters could be used to modify the diameter of the

nanotubes and thereby control the rate of diffusion of the drug stored in them into the

biological environment.

309

Surface texture of the material is also a property of interest; for example, electropolishing of

a Ti-6Al-7Nb alloy decreased the amount of S. aureus adhering to it.

310

A tradeoff,

however, is expected to arise because osteoblasts, which compete with the bacteria for the

bioactive surface of the implant in a process that greatly determines the clinical outcome,

311

also prefer to attach to rougher surfaces, such as those that typify naturally topographically

irregular calcium phosphates.

312

To that end, titanium implants are being subjected to

sandblasting and etching procedures,

313,314

as well as coated with bioactive layers,

predominantly hydroxyapatite,

315–317

to make up for their intrinsic bioinertness and have

their bioactivity boosted before surgical insertion.

G. Composites

In the design of nanoparticles for biomedical applications great emphasis has been placed on

particles capable of simultaneously aiding in prevention, early detection, and treatment of a

medical condition. With antibiotic calcium sulfate cements already in use in prophylaxis

against surgical wound infection, it can be expected that theranostic particles able to

simultaneously prevent, monitor, or diagnose the onset of infection and release antimicrobial

agents to prevent its early spread may be developed in the future. In that sense fantastic

multifunctional composite nanoparticles can be considered to be the ideal toward which the

nanoparticle fabrication field will advance (Fig. 7a). The difficulties in achieving stable,

chemical functionalization of calcium phosphate particles with therapeutic ligands can be

mitigated by coating them with a chemically bondable layer, such as PLGA

318,319

PCL

320

or by forming around the calcium phosphate core multilayered composite particle

structures

321

with an ability to carry various therapeutic agents either between the calcium

phosphate layers or within the polymeric coatings (Fig. 5b–d). Through a simple series of

chemical steps, polymeric coatings can also be conjugated with various targeting or

therapeutic ligands.

322

Binding amino acids with appropriate physical properties via PEG

linkers, for example, phenylalanine as a hydrophobic residue, lysine as a positively charged

one, and glutamic acid as a negatively charged one, can be used to increase the drug-binding

affinity of the polymeric surface.

323

Tailoring of the nanodiamond particle surface with

carboxylic or amino groups to render it negatively or positively charged under physiological

conditions, respectively, greatly affected the binding and release of various drugs; binding of

the negatively charged drug by physisorption to an amine-functionalized surface is so

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intensive that virtually no release in vitro occurred.

324

Such combinations of various loading

locations on the particle could potentially yield multiple-stage release profiles that might not

only favor the antimicrobial efficacy of the particles in vivo but also prove to be beneficial in

increasing the regenerative capacity of the carriers, given that the tissue regeneration process

following injury can be divided into multiple stages (Fig. 7b), each of which could be

targeted and augmented by a specific particle additive released within a precisely tailored

time window. To avoid adverse outcomes resulting from obviation or incompletion of any

single one of the interconnected steps in the bone-healing cascade (Fig. 7c, d), the

biomolecular machinery involved in every one of these stages could be targeted separately

and triggered at the right time by using such a smart composite particle that sequentially

releases its multiple payload in a highly controlled, spatiotemporal manner. An interesting

approach to achieving such multimodal release profiles is through cooperative assembly of

block copolymers as elemental building blocks of the particle, each of which carries a

unique therapeutic payload and degrades at a different rate.

325

The combination of calcium phosphates with a polymeric component can also be beneficial

for the second essential function to be achieved by these nanoparticulate drug carriers, in

addition to their antibacterial role: assistance in bone regeneration. Namely, since bone itself

is a composite material comprising a soft, collagenous component and a hard, ceramic one,

it is natural to expect that a soft/hard composite of a similar nature should prove an ideal

material for bone replacement therapies. In view of this, a range of properties of calcium

phosphates is improved upon their combination with a polymeric phase, starting, most

essentially, with the mechanical ones. Namely, it is generally assumed that the

microstructure and nanoarchitecture of calcium phosphates alone cannot be modified in such

a manner as to make the material mechanically compatible with the grafted bone and

prevent the frequent fracture of the filler upon its surgical placement to substitute natural

bone.

326

Only a combination with a soft component is thought to be able to ameliorate these

fundamental issues associated with the clinical application of calcium phosphates. The

combination of viscoelastic properties of the polymers and osteoconductivity of calcium

phosphates has yielded composites that surpassed the resistance to fracture, structural

integrity, and stiffness of the individual components,

327

making up for the low compressive

strength of the former and the brittleness and lack of malleability of the latter.

328

Reinforcement with polypropylene fumarate,

329

for example, improved the flexural strength

of brushite from 1.8 to 16.1 MPa and increased the fracture surface energy from 2.7 to 249

J/m

. Although calcium phosphates exhibit relatively high values of compressive strength

(10–100 MPa), as opposed to tensile and shear strengths (1–10 MPa), even these values

could be improved with the addition of a polymer, as exemplified by the doubling of the

compressive strength of a biphasic calcium phosphate cement, from 35 to 60 MPa, upon the

incorporation of only 0.5 vol% of a superplasticizer based on a vinyl-modified

copolymer,

330

as well as upon the addition of gelatin

331

or ammonium polyacrylate.

332

Polymers could also increase the plastic flow and enhance the viscosity of the material, thus

making possible its preparation in the form of an injectable self-setting paste,

333

although

there is usually a fine line dividing an excessive increase in the setting time from improved

mechanical properties.

334,335

Finally, the resorption time and the corresponding bone

ingrowth rate significantly increased when hydroxyapatite was implanted as a bone

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substitute in vivo in a composite form, in combination with PLGA,

336,337

a polymer that is

able to accelerate the resorption of calcium phosphates by releasing its acidic degradation

products.

338

Yet another tradeoff exists here: An increase in the porosity of the ceramic

substructure of the composite leads to improved bioresorption characteristics but

simultaneously entails an increased susceptibility of the material to crack propagation and

the corresponding proneness to fail under load-bearing conditions.

339

Such composite particles showed promise in earlier research. Gentamicin-containing

granules composed of hydroxyapatite nanoparticles, chitosan, and ethyl cellulose, for

example, were effective in the treatment of chronic osteomyelitis.

340

Another prospective

hybrid organic–inorganic system was formed by dispersing silsesquioxane microspheres

loaded with acetylsalicylic acid as an anti-inflammatory model drug in a calcium phosphate

cement.

341

Similar composites reduced in size to the nano scale may be recognized as a

trend toward which this field will be moving (Fig. 8). Polymeric coatings may also increase

the loading capacity and prevent the burst release of the drug merely adsorbed on the

particle surface. Coating chitosan/tricalcium phosphate composites with 2.5w/v% PCL has

thus mitigated the burst release effect and promoted zero-order kinetics for the release of

vancomycin during the first 6 weeks.

342

Impregnation of the poly(α-hydroxy esters) with

bone morphogenetic proteins has been shown to (1) overcome the inflammatory response,

(2) induce full bioresorption of the polymer, and (3) enhance bone growth,

343–345

while the

addition of demineralized bone particles to PLGA reduced (1) inflammation, (2) fibrous

tissue encapsulation, and (3) foreign body giant cell response.

346

The combinations of

alkaline calcium phosphate phases, such as hydroxyapatite or octacalcium phosphate, with

acidic poly(α-hydroxy esters) are thus particularly interesting because of their ability to

mutually compensate for potentially harmful pH changes that follow their degradation. The

wide range of pH conditions provided by the synergetic action of osteoblasts and osteoclasts

in the degradation of calcium phosphates in vivo makes the use of pH-sensitive coatings

potentially interesting, too. Poly(aspartic acid) presents one such pH-sensitive polymer; its

swelling is more pronounced at the physiological pH than at pH ~3 and can be facilely

controlled by the degree of cross-linking.

347

Another type of environmentally responsive

polymers are thermosensitive polymers, which transform from sols to hydrogels at body

temperature and enable in situ gelling at the target site promptly after injection.

348

Some of

the biodegradable polymers of this type include N-isopropylacrylamide copolymers,

poly(ethylene oxide)/poly(propylene oxide) block copolymers, and PEG/poly(D,L-lactide-

co-glycolide) block copolymers, the latter of which have been successfully applied to

encapsulate teicoplanin with 100% efficacy and treat osteomyelitis in rabbits.

349

That even the activity of antibiotics can be improved with a proper coating is illustrated by

the more effective prevention of the formation of S. aureus biofilms when vancomycin was

delivered encapsulated within cationic liposomes and carried in a porous nano-

hydroxyapatite/chitosan/konjac glucomannan scaffold.

350

The same antibiofilm effect was

achieved by the delivery of liposomal gentamicin from scaffolds containing β-tricalcium

phosphate with release kinetics able to be controlled by the liposome size.

351

Functionalization of particulate carriers with anionic amphiphiles that may disrupt the

bacterial biofilm and neutralize the carbohydrates by the action of which bacteria penetrates

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the cell membrane is another unexplored but potentially fruitful direction for research. With

antioxidant therapy being one of the hotspots of medicine, endowing carriers with reducing

agents, such as ceria domains

352

or ascorbic acid,

353

could be an avenue for abating the

reactive oxygen species and minimizing the oxidative stress that entail infectious disease.

Nanoparticle uptake by the cells can be controlled using ζ potential,

354

a ubiquitous physical

property,

355

but precise correlations between the surface charge and the therapeutic efficacy

of nanoparticles in the treatment of osteomyelitis have yet to be established, even though

wound healing could be enhanced by endowing cells with relatively high ζ potentials.

356

The usage of dispersion agents or the application of other strategies from the repertoire of

colloid chemistry

357

to promote greater dispersion and penetration of the antibiotic-carrying

particle to the infected tissue—a general challenge for the developers of injectable drug

delivery materials

358

—is another unexplored avenue.

Conjugation of the carrier particles to moieties that would have an affinity for various bone

components

359

and act as either targeting agents or metabologens is yet another unexplored

research directive in the design of antiosteomyelitis composite particles. Human

recombinant BMPs, two of which—BMP-2 and BMP-7—were approved for specific

clinical cases by the FDA, have been successfully delivered using various types of

nanoparticles, ranging from poly(α-hydroxy esters) to PEG-based hydrogels to dextran to

polymeric composites with hydroxyapatite to calcium phosphates alone,

360

and their

covalent binding on the polymeric particle surface may prove to be a more effective

approach for their delivery compared with internal encapsulation, especially in view of the

extraordinary sensitivity of their osteoinductive effect to the release kinetics.

361

These

conjugates could also include biomolecules that inhibit specific bacterial ingredients, such as

(1) lipoteichoic acids, components of gram-positive cell walls that induce bone resorption;

(2) polysaccharides in the bacterial capsules, which play a role in the adhesion of bacteria

onto an osseous or implant surface and the formation of a biofilm, the basis for proliferation

of pathogens in hard tissues; or (3) other osteolytic factors, including cytokines or other

signaling molecules, which may interfere with the pathway of the osteoblast lineage.

364

Such efforts may have a chance to bring researchers from drug delivery and drug discovery

fields closer because, after all, the synergy between the drug and the particle will prove to be

of ever more vital importance in the design of ultrapotent therapeutic agents in general.

Inclusion of peptides with strong antibacterial properties, which tend to be more immune to

promoting bacterial resistance if delivered in concentrations lower than minimal inhibitory

ones, would present another interesting approach.

365,366

The polymeric surface of a

composite particle could be functionalization with arginylglycylaspartic acid, a tripeptide

involved in cellular recognition and capable of triggerinxg adhesion of fibroblasts.

367

Bisphosphonates, molecules with a strong affinity for the mineral component of bone

368,369

and most commonly prescribed in the prevention and treatment of osteoporosis and other

conditions featuring bone loss and fragility,

370

could be used further to ensure particle

localization and the delivery of therapeutics directly in the area of infected tissue. Although

a possible concern comes from the clinically observed adverse consequences of

oversuppressed bone resorption and disrupted bone metabolism by the prolonged use of

bisphosphonates,

371

the risk for developing these side effects is still small compared to the

benefits.

372

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ZnO has been added to implantable PMMA beads as a radiographic contrast medium for the

past 30 years,

373

yet calcium phosphates and other carriers could be doped with heavy metal

atoms such as

111

In,

99m

Tc, Gd, or Mn or bound to optically active molecules and used for

the same imaging purpose with far greater sensitivity. Combinations of calcium phosphate,

PLGA, and semiconductor quantum dots

374,375

at the nanoparticle scale have enabled

monitoring of the particle route in the body, and the distribution of the locally implanted

therapeutics could be monitored in a similar manner. Quantum dots are, however, known for

their cytotoxic nature,

376

with only a few exceptions, including silica-based compositions,

the only type currently approved for use in clinical trials by the FDA.

377

Proposed as

bioimaging alternatives to inherently toxic quantum dots and nonbiodegradable aromatic

polymers are aliphatic, biodegradable, and tunably photoluminescent oligomers,

378

but they

have yet to be explored as components of bone tissue substitutes.

Porosity of composites in the compact, fully set form could be modified using other

additives, such as glucose,

379

calcium sulfate,

380

calcite,

381

gelatin,

382

or others, and set to a

specific pore size, pore size distribution, and pore interconnectivity that maximize the

internal cell proliferation and the transfer of nutrients and metabolic products. For example,

combinations of silica and calcium phosphate allowed for a control over porosity of the

resulting gentamicin-loaded nanocomposites in the mesoporous (2–50 nm) and macroporous

(>50 nm) ranges by means of controlling their silica content.

383

Porosity also could be

limited to the surface only

384

to promote a bioactive response while preserving the

compactness and stability of the core of the system against attack from the corrosive

biological environment, or the other way around, porous on the inside and compact on the

outside,

385

like bone itself. It could be also made gradient, extending throughout the bulk of

the composite in different ways.

386

As a matter of fact, Janus-faced

387

and functionally

gradient structuring

388

on the nano and molecular scales are other largely unexplored, yet

incredibly potent features of the next generation of advanced materials. Finally, enriching

antibiotic carriers with pluripotent cells, such as mesenchymal stem cells able to

differentiate into osteoblasts,

389,390

would be another research step in the direction of

advanced therapeutic platforms for simultaneous bactericidal and osteogenic performance.

In that sense, understanding the role of the extracellular matrix and an array of

microenvironmental cues in directing the pluripotent cell fate currently stands as a major

challenge to be overcome to minimize cells’ chances of acquiring neoplasticity and to

maximize their chances of acquiring the best possible phenotype for the given therapy.

391

V. ADDITIONAL CHALLENGES

A. Inconsistencies Arising from Different Analytical Contexts

An essential wisdom conveyed from the drug delivery field is that context is everything.

When not delivered in a proper manner, even the most effective therapeutics will be

deprived of their remedial effectiveness. A direct corollary of this insight is that drug

discovery and drug delivery could be imagined as two sides of the same coin,

complementing each other in a complete drug therapy. In other words, in a wrong setting

even the most therapeutically potent agent will be ineffective, whereas even the most toxic

chemicals applied in the right amount and setting could strengthen an organism, as data in

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support of the effect of hormesis could indicate. Therefore, it should not be surprising that

problems and challenges will continue to abound even if we were to succeed in the design of

a perfect drug delivery carrier.

An example of the effect of the environmental context on a drug elution profile from a

particulate carrier is shown in Fig. 9. Namely, two kinetically distinct release profiles from

an identical drug-containing powder result, depending on the release measurement method

applied: zero-order when a comparatively small volume of the solvent is replenished daily

and first-order when a larger volume of the solvent is used without its daily replenishments.

Whereas the drug is released with a burst effect when large amounts of solvent surround the

powder, a limited volume of the solvent limits the maximal amount of the drug that could be

released before saturation, leading to identical concentrations of the drug in the solution

sampled at regular time spans (24 hours). In such a manner reprecipitation of sparsely

soluble hydroxyapatite on the surface of more soluble di- and tricalcium phosphates in the

form of a protective layer that hinders further dissolution, which might have occurred in vivo

or at smaller solvent volumes, and higher corresponding degrees of supersaturation can be

prevented by frequently replenishing the solvent and used to speed up the dissolution of

these calcium phosphate phases under physiological conditions. Neither of these two

methods, however, mimics the biological conditions under which fast clearance of the

released drug is typically observed, nor do they account for the effects of the complex

interface between the device and various macromolecules and cells of the host organism on

the drug release. Although the size, shape, and elasticity of nanoparticles in biological

milieus do influence their biodistribution profiles, the route of uptake and the mechanism of

interaction with a cell are mainly determined by the protein corona adsorbed on the particles

and the surface propensities that it endows them with.

394,395

How to devise in vitro drug

release testing procedures that would be able to replicate in vivo conditions better, typically

characterized by (1) a more dynamic flow of fluids, (2) specific local pH profiles that are

often disease-dependent, and (3) much more complex and selective media, is another

colossal challenge for the drug delivery field.

Incompatibility between in vitro and in vivo tests has been frequently observed

396

; what has

been shown as toxic or inflammatory in vitro can have the same effect in vivo but can also

provide the right level of inflammation that follows every successful reaction toward unity

between the bone and the implant.

397

The trivial observation that pure water momentarily

destroys cells in culture via osmotic rupture of the cell membrane, whereas we consume it

every day without serious consequences, could be used to illustrate the inevitable

discrepancy between testing materials in culture and in a more complex, organismic

environment. Moreover, not only do different lines of the same cell type often respond

differently to identical chemical stimuli

398

; the same insight applies to cells from the same

line but at different stages of cell cycle progression.

399

On the other hand, techniques for

replicating the exact microenvironment in which cocultured primary cells exist in the body,

which is one of the key factors that determine their fate,

400

have yet to be developed.

As for animal models of osteomyelitis, different species and experimental protocols have

been used in the past. The rabbit is the oldest animal model of osteomyelitis, dating back to

1941,

401

and it has traditionally involved injections of a suspension comprising S. aureus

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and 5 wt% sodium morrhuate, a sclerosing agent causing aseptic necrosis of bone, to induce

suppuration and subperiosteal abscesses in the bone. Fibrin glue and other sealants often are

used to prevent bacterial leakage.

402

The most common, however, is the rat model, typically

involving the insertion of screws or other fixation devices inoculated with S. aureus.

403

alternative method involves the creation of a defect in the tibia or femur by drilling a hole in

it and stabilizing it with screws, plates, and wires. The defect then is filled with collagen or

gelatin sponges soaked in S. aureus and kept there for up to 24 hours, which is superseded

by injection or implantation of an antibiotic-releasing material.

404

Mice and chickens are

other common small-animal models, whereas dogs, goats, and sheep are the most common

large-animal models, whose main advantage is the ability to accommodate real-size devices

and tolerate multiple interventions, alongside more veritably mimicking the mechanical

loads born by bones in the human body. A central challenge in all of these models is how to

create bone lysis that is sufficient but not excessive and does not threaten the fixation

stability. Avoiding the formation of virtually untreatable biofilm on fixation devices is cited

as another challenge faced by these animal models,

405

even though it presents a more

faithful model for cases in which infection is secondary to surgical foreign body

implantation. Yet another detail common to these models is that chemically induced necrosis

or mechanical trauma, such as fracture (as in most mouse models) or the insertion of

intramedullary pins, plates, or other fixation devices, need to be coupled to dispersal of the

pathogen to induce chronic infection.

406

Still, in spite of more than 70 years since the first

reproducible animal model of chronic osteomyelitis was reported and continuous

advancements since then, inconsistent correlations between in vitro and in vivo antibacterial

efficacies of therapeutic agents still commonly occur.

407

As for the drug elution rate, values obtained in vitro may drastically increase in vivo for

various reasons, including the complex interface with chemical and biological species (e.g.,

enzymatic activity, the concentration of free radicals that induce oxidative scission of

covalent bonds between monomers, the types of antibodies adsorbed, or the extent of fibrous

capsule formation, if any) or different rheological properties, demanding new strategies to

ensure the optimal 4–6 weeks of release time. PLGA scaffolds, for instance, degraded faster

in vivo than in vitro,

408

and in addition to phagocytosis, enzymatic hydrolysis, the regions of

low pH at the cell–material interface, and biomechanical stress, increased wetting in

biological conditions can be another factor responsible for this effect.

409

This is especially

relevant for hydrophobic polymers, the category to which unmodified PLGA belongs. This

disparity between the carrier degradation and the drug release kinetics estimated using in

vitro and in vivo measurement modes is expected to be even higher for calcium phosphates

than for polymers because while the degradation of the latter is primarily driven by

hydration and hydrolysis both in vivo and in vitro, the degradation of ceramic implants is

mainly caused by the phagocytic and acidifying (pH 7.4 → 3–4)

410

action of multinucleated

osteoclasts and macrophages.

411

This discrepancy is further added up to by knowing that

calcium phosphate phases more soluble than octacalcium phosphate transform to a certain

degree into the most stable phase, hydroxyapatite under physiological conditions.

412,413

Factors that dictate to what degree this transformation takes place in vivo are not clearly

defined; on one hand its precursors are classified as biosoluble ceramics,

414

unlike

bioresorbable hydroxyapatite, and indisputably degrade faster than the latter, whereas on the

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other hand this transformation has been both theoretically predicted and experimentally

verified on numerous occasions.

415–417

Comparatively soluble calcium phosphates,

including monetite, brushite, and amorphous calcium phosphate, were also found in vivo,

418

suggesting that this transformation may frequently be limited to a few surface layers that

undergo intense dissolution/reprecipitation, thus protecting the inherently unstable and

soluble particle or implant core. The chemical propensity of more soluble calcium

phosphates to be resorbed at a higher rate, however, is complicated by knowing that

osteoclasts anchor more steadily on the less soluble calcium phosphate phases; the

substantial release of Ca

ions from the surface of the more soluble ones disrupts the

ordering of actin microfilaments in the osteoclast podosomes, leading to the periodic

detachment of the cells from the material surface.

419

Similarly, whereas the degradation of

aliphatic polyesters(e.g., PLGA)is slowed under fluid flow because of dissipation of acidic

byproducts that would have sped up the degradation process, fluid flow prevents local

supersaturation and accelerates the degradation of alkaline calcium phosphates such as

hydroxyapatite. Hence, despite the excellent release profiles in vitro, the possibility that the

minimum inhibitory concentration in sections of the target tissue may not be exceeded in a

time-sustained manner, be it due to biofilm formation, intracellular colonization, premature

biodegradation, or other factors, will always exist. In such a way exist threats that antibiotic

resistance could be inadvertently promoted, an effect that directly contributes to the global

loss of antibiotic efficacy in use.

420

In those cases even a direct injection of a bolus dose of

the antibiotic may lead to more favorable outcomes than the implantation of the drug–carrier

composite.

421

B. Antibiotic Specificities

Antibiotics differ according to their mechanism of action; some, such as β-lactam

antimicrobials, are time-dependent, requiring prolonged presence in the target zone for

effective suppression or eradication of the pathogen, whereas others, such as quinolones and

aminoglycosides, are concentration-dependent, requiring higher concentrations over shorter

periods of time.

422

Therefore, depending on the nature of the drug, differently structured

carriers may prove to be most optimal. As expected from the synergetic background of drug

carrier interaction, the properties of the carrier influence the efficacy of the drug therapy, but

the drug identity, amount, and binding mechanism in turn influence the properties of the

carrier. The anionic group of gentamicin sulfate, for example, had an inhibitory effect on the

crystallization of brushite during coprecipitation of the antibiotic and the carrier, resulting in

smaller particles, lower porosity, and slower drug release compared with pure

gentamicin.

423

The morphology, crystallinity, and dispersability of particles coprecipitated

with the drug may thus all be affected by the drug properties. The loading efficiency and the

release rate of a drug from a particle is consequently dependent on the molecular nature of

the drug, as exemplified by the different elution profiles for different antibiotics

424

The

release from PMMA beads, for example, varied drastically, reaching completion anywhere

between 3 and 21 days, depending on the antibiotic admixed.

425

Also, whereas the release of

vancomycin from a brushite cement reached completion between days 1 and 2, only one

quarter of tetracycline loaded within the same carrier was released on day 5.

426

Synergetic

effects are important here; for example, the elution rates for tobramycin and vancomycin

released together from a PMMA cement were higher than those when the drugs were

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released alone.

427

Then, not only does the identity of the drug have an effect on release

profiles, but its amount can drastically modify them, too, as exemplified the case where 25

μg/cm

of paclitaxel deposited on the surface of a phosphonohexadecanoic acid coated

cobalt chromium alloy released 90% of the drug in the first week, whereas quadrupling the

paclitaxel concentration to 100 μg/cm

resulted in zero-order release throughout a 5-week

period of time.

428

The release zone of the antibiotic following surgical insertion is generally only a few

centimeters, and the rheology of the fluid in which the drug carrier is immersed, including

hematoma and seroma, might redirect the released drug away from the target site. Therefore,

continuous-flow chambers have been designed to assess drug elution profiles under more

dynamic conditions that resemble the in vivo context to a greater extent.

429

Still, different

implantation sites greatly affect the biodegradation rate of the implant, including the release

profile of the drug that it contains.

430

Subcutaneous implantations of a biomaterial

composed of PEG and poly(butylene terephthalate) thus degraded faster than the

intramuscular ones,

431

whereas polydioxanone orthopedic pins were resorbed faster when

implanted in the medullar canal rather than intramuscularly or subcutaneously.

432

Mechanical loading, sheer, and friction are other factors that contribute to the different

release from orthopedic drug delivery devices implanted at different sites in the body.

433

Also, hypochlorites generated smaller polymeric fragments with higher toxicity than

peroxides,

434

suggesting that the dominant reactive oxygen species as inflammatory

compounds in the implantation area, which are involved in the degradation of polymeric

carriers, inevitably define the toxic propensities of the biomaterial in question. In that sense

the art of surgical implantation has to complement the reliability of the drug release pattern

of the implanted drug–carrier composite. That different target areas in the body require

unique drug delivery platforms for optimal release has been confirmed many times, and it is

logical to expect that the same will prove to be true in the treatment of different segments of

bone. Mandibular infection, for example, typically requires a shorter duration of antibiotic

therapy compared with long-bone infection. Intensely vascular cancellous bone, with a

comparatively high rate of turnover, may thus be expected to require more intense release

kinetics compared with less vascular and more slowly remodeled cortical bone.

435

general, the principle similia similibus curantur, dictating the substitution of like with like,

is expected to apply in every aspect of tissue engineering, including the province of bone.

C. Intracellular Colonization by S. aureus

Another potential difficulty arises from the fact that S. aureus, the main causative agent of

osteomyelitis, found in healthy oral and nasal flora has the ability to penetrate endothelial,

epithelial, and osteoblastic cells and thrive in the intracellular environment, where it is less

susceptible to the antibiotic therapy.

436–438

Different but in all cases finite uptake efficiency

and kinetics were observed for different strains of S. aureus, and common to all of them was

the role of fibronectin-binding proteins in intracellular colonization, the blocking of which

completely prevented the latter from occurring.

439

S. aureus internalized by the cells and

shielded from the host immune system is thought to provide a reservoir of bacteria in

recurring osteomyelitis and inhibit the immunological role of osteoblasts in releasing

cytokines and attracting leukocytes to the infection site. Targeting the antibiotic therapy to

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these intracellular colonies may thus prove to be more relevant for treating chronic bone

infection than eliminating only pathogens that colonize the bone matrix.

440

Nanoparticles,

including calcium phosphate ones, have shown great promise in the intracellular delivery of

plasmids,

441–443

and they may similarly prove to be effective carriers of antibiotics inside

the cell where bacterial microcolonies are localized. Correspondingly, a recent study showed

that clindamycin-loaded hydroxyapatite and amorphous calcium phosphate particles are

more effective in reducing the intracellular bacterial population and slowing the growth of S.

aureus cocultured with osteoblastic MC3T3-E1 cells than the pure antibiotic.

444

Figure 10a

shows a schematic description of their uptake by the cell and the delivery of a genetic

material to it, whereas 10b shows the intracellular localization of calcium phosphate

nanoparticles in an immunofluorescent analysis of osteoblastic MC3T3-E1–calcium

phosphate interface. A reduction in the number of intracellular bacteria has already been

demonstrated for nanosized PLGA particles loaded with nafcillin.

445

The rapid degradation

of calcium phosphate carriers in the acidic milieu of a lysosome–endosome complex

following uptake may also increase the osmotic pressure and enable the plasmid or protein

cargo to escape swiftly into the cytoplasm before it is enzymatically hydrolyzed, which

could be considered yet another advantage of calcium phosphates as intracellular delivery

agents.

D. Synergetic and Sensitivity Effects

Immunofluorescent labeling and confocal microscopy, along with other in vitro assays,

including real-time polymerase chain reaction, can provide good insight into the cell–

material interface on which the bone regeneration aspect of therapy ultimately depends

446

(Fig. 11a). Nanoparticles have, however, been notorious in terms of resisting any clear-cut

descriptions of their biological effects, as exemplified by the recently derived inverse, dose-

dependent toxicity relationship for 100-nm silica nanoparticles in the human epithelial

intestinal HT-29 cell line.

447

In a similarly counterintuitive fashion, the uptake efficiency

and the expression of plasmids could occasionally be out of proportion: Efficient uptake

may lead to low gene transfection and vice versa.

449

Synergetic effects resulting from the

minor amounts of impurities left over from synthesis procedures or supposedly inert product

components may prove to be equally important in determining therapeutic outcomes. For

example, the osteogenic effect of calcium phosphate as a drug carrier was able not only to

mitigate but also to fully reverse the unviable effect that the pure antibiotic exerted on

osteoblastic cells, while retaining its antimicrobial potency through a more sustained release

of the antibiotic (Fig. 11b).The addition of calcium hydroxide to tobramycin-containing

PMMA beads similarly had a protective effect on bone against the high concentrations of

the antibiotic.

450

A related problem occurring in parallel with the sophistication of the carrier particle is the

difficulty transferring the synthesis methods from the laboratory to the clinic or any large-

scale fabrication setting. Namely, the more intricate the particle, the greater the range of

experimental variables to which its preparation is sensitive.

451

The case of Abbott Labs

losing hundreds of millions of dollars trying to restore the polymorph of their proprietary

AIDS drug, Ritonavir, after a new polymorph had begun to appear in their synthesis batches

and before being forced to withdraw the drug from the market and lose another half a billion

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dollars,

452

can thus be instructive, if not ominous, in its foreshadowing the trend of a

pervasive irreproducibility toward which the science of synthesis of fine particles streams.

Startups in organic photovoltaics have, for instance, failed to make profit mainly because of

the unsurpassable variability in processing outcomes. A systematic analysis of two donor

materials synthesized by the same manufacturer resulted in confirmation of the same

crystallinity, side-chain variation, interface composition, and domain morphology, yet their

performance was radically different because of a difference in the nano scale that was too

fine to be probed by state-of-the-art instrumentation.

454

Note also that the purity of solvents

and compounds used as precursors for the synthesis of nanostructured powders has had a

drastic effect on their morphology,

455

whereas cell detachment or excellent spreading result

depending on whether cultured cells are grown on ordinary or tissue culture–grade

polystyrene.

456

Alongside the chemical effects exemplified by different reaction outcomes

when reactants produced by different manufacturers are used, there exist examples of the

following physical effects on synthesis reactions: stirring rate

457,458

; the Earth’s magnetic

field

459,460

; gravity,

461,462

whose intensity appears to be directly proportional to the

osteogenic potency of mesenchymal stem cells and boney tissues

463–465

; the seasonal

variations in (1) humidity,

466

(2) barometric pressure, and (3) temperature; micrometric

differences in the positioning of samples in furnaces during annealing

467;

experimental

animal behavior; and circadian rhythm, on which the expression of almost 50% of genes in

certain cell types has been shown to depend

468

; and the reaction vessel composition,

469

texture,

470

and dimensions,

471,472

one of the most critical parameters in the transfer of

synthesis methods from a small-scale setup in a laboratory to a large-scale fabrication setting

in an industrial milieu. For example, overly tall reaction volumes increased the probability

of the formation of aragonite or vaterite, two of the less thermodynamically stable calcium

carbonate phases during the precipitation of this compound from a solution in the presence

of an organic matrix, whereas flattened volumes were more prone to yield calcite, the least

soluble calcium carbonate phase.

473

Agitation, a traditional means for dispersing particles,

has occasionally had the opposite effect, inducing the aggregation of both polymeric and

inorganic nanoparticles,

474

which may explain cases in which the efficiency of drug loading

via adsorption is inversely proportional to the stirring rate.

475

Similarly antagonistic effects

of specific physicochemical synthesis parameters present more of a rule than an exception in

the field of nanoscience.

476

Assessments of biological responses are particularly prone to

exhibit such antagonistic intricacies, as exemplified by the case in which doubling the dose

of calcium phosphate nanoparticles transformed the response of human bronchial epithelial

cells from unviable to viable.

477

Biological systems in general are, in fact, notorious for

their sensitivity to the slightest change in their homeostatic equilibria, as demonstrated by

the long-familiar finding that bone reaches its maximum toughness in compositions that

leave behind 66.5 wt% of ash, twice higher than the value at 65 and 68 wt% of ash

content.

478,479

On the characterization side, this sensitivity results in ever more difficult

separation of the effects of the measurement system from the properties of the measured

objects, and this will increasingly apply to all 3 essential characterization aspects toward

which materials science progresses: single-particle spectroscopic microscopy, high-

throughput analysis, and in situ analyses. Of course, each of these problems conceals a

gateway to an exciting opportunity, as exemplified by the fact that ultrafine microscopic

methods do regularly create or remove single vacancies

480

or shift adatoms

481,482

on the

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surfaces of analyzed materials, but measuring the energy required for these local

transformations to take place has allowed for the construction of atomic-scale images of

local phase properties,

483,484

which is state-of-the-art in the field of materials imaging.

In view of this one thing is certain: the design of a new generation of biomaterials will need

to cope with the very same issues pertaining to the extraordinary sensitivity of function to

the finest structural and compositional variations if its products are to match perfectly the

properties of the natural tissues they are meant to replace. Yet, as common wisdom has it, a

path without risks, perplexities, and challenges is not worth taking, and difficulties arising

from attempts to design nanoscale platforms for the ultrasensitive delivery of pharmaceutics

can certainly be seen as signs that those involved in these endeavors are heading in the right

direction.

VI. SUMMARY

We approach a time of a prolific confluence of materials science and medicine. It is

anticipated that materials science will provide the foundations for the design and

development of advanced diagnostic and therapeutic methods. Within the frame of its

objective, this review has provided a critical view of the current state of affairs in the

development of nanoparticulate and other solid-state carriers for the local delivery of

antibiotics in the treatment of osteomyelitis. In a broader picture extending outside of this

narrow frame, however, is a view of ongoing progress in the way a relatively modest field of

medicine, in terms of the complexity of therapeutic materials used, is being revolutionized

by recent advancements in materials science and engineering. The surgical implantation of

PMMA beads or, occasionally, plaster of Paris still presents the most popular method for the

local and sustained delivery of antibiotics in the treatment of osteomyelitis, but this is about

to change as more sophisticated materials, nanostructured in essence, are being developed

on laboratory benches and sent down the translational path toward the direction of the

bedside.

The two principal downsides of the traditional means of treating bone infection are systemic

and long-term administration of antibiotics and the necessity for surgical debridement. The

new generation of carriers for the delivery of antimicrobials envisaged during this discourse

is expected to tackle these issues by first promoting the sustained release of antibiotics

limited to the target site. Another vital feature of advanced carriers for the controlled release

of antibiotics is their ability to contribute to the osteogenesis of the adjacent tissue in parallel

with their degradation and replacement with regenerated bone. Thus antibiotic delivery and

tissue regeneration are two central aspects of therapies for osteomyelitis in which milestone

improvements are to be expected using the new generation of nanostructured carriers. Next

steps in their development would be tuning their structure to an environmentally responsive

and spatiotemporally targeted performance, as well as integrating them with alternatives to

traditional antibiotics, whose ineffectiveness against increasingly resistant opportunistic

pathogens is approaching critical scales. If we look at this trend in the development of drug

delivery carriers for treating a particular disease from a broader angle, outside of the frame

once again, we might conclude that structurally complex, multifunctional, theranostic

composite nanoparticles present an object of interest toward which scientific efforts

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colliding at this multidisciplinary junction of medical significance will be converging in the

years and perhaps decades to come.

For a long time the degradation of polymeric biomaterials had been considered an

unfavorable process, a cause of the deterioration of their properties and performance in the

body. It took a fresh, new look at them to turn these demerits into advantages and present

their degradation as a colossal value and untapped potential for the biomedical community.

Using their degradation as a pathway to controlled drug release and tuning it to the rate of

new tissue in growth nowadays presents the basic approach to tissue engineering and

regenerative medicine. It goes without saying, of course, that the time for the reversal of this

paradigm and inauguration of equally legitimate strivings to create bionic tissue-engineered

materials—whose purpose would be not only to restore the lost functionality in a segment of

the body or the body as a whole but also to augment and raise it far beyond the levels of

ordinariness via lasting assimilation of the new interfaces—has yet to come. Be that as it

may, potential problems faced by even the hypothetically perfect antibiotic delivery vehicles

mentioned toward the end of this discourse include (1) the propensity of S. aureus, the main

causative agent of osteomyelitis, to form intracellular colonies involved in recurrent, chronic

osteomyelitis; (2) the need for the mechanical and release properties of a carrier to be

adjusted to the target area of surgical implantation or injection; (3) the disparity between

environments in which in vitro and in vivo drug–carrier composite testing is carried out; (4)

unpredictable synergetic effects of the delivery system components or foreign agents; and

(5) experimental sensitivity issues posed in parallel with the increasing subtleness of

nanoplatforms designed for the controlled delivery of therapeutics. Inspired by the parable

of biodegradable polymers, we could conclude that all of these problems, were they

considered from a fresh, new angle, might either stimulate the development of

bioengineered systems that will solve them and many other problems at the same time or,

even more amazingly, be glimpsed as solutions per se to problems existing in a different

domain. For if a new, multidisciplinary model for a prolific lifetime in science teaches us

something, it is that two gaps in knowledge, when combined, can create a bridge to much

greater knowledge lying far beyond the horizon.

Acknowledgments

Writing of this review was supported by the National Institutes of Health grant R00-DE021416: Osteogenic

calcium phosphate nanoparticles with designable drug release kinetics. The confocal optical micrographs presented

herein were acquired by the author at the Nikon Imaging Center at the University of California, San Francisco.

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FIG. 1.

(a) Age of patients with osteomyelitis.

(b) US population ≥65 years old (bars), along with

a projected increase in the number of patients with bone disease. Sources: US Bureau of the

Census and Office of the Surgeon General.

Uskoković Page 53

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 2.

The number of prosthetic joint infections (solid circles and squares), increasing in direct

proportion with the total number of knee (squares) and hip arthroplasties (circles)

performed.

Uskoković Page 54

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 3.

Radiographs of tibiae displaying a clinical case of acute pyogenic hematogenous

osteomyelitis, also known as Brodie’s abscess (△), along with an area of increased bone

density around the lytic lesion due to periosteal reaction and osteosclerosis (□) (a), and

evidence of chronic osteomyelitis from a goat model, manifesting as a periosteal reaction in

the proximal area of the bone (b).

Uskoković Page 55

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 4.

A typical preoperative preparation of poly(methyl methacrylate) (PMMA) beads

encapsulating an antibiotic of choice, consisting of (1) manual admixing of the antibiotic

with the bone filler; (2) filling a mold with the solid mixture; (3) hardening; (4) removing

the beads; (5) collecting the beads.

Uskoković Page 56

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FIG. 5.

(a) Narrowly dispersed calcium phosphate nanoparticles prepared by precipitation form

aggregates upon desiccation at low pressure, capturing the antibiotic clindamycin inside of

the resulting pellet pores, thus ensuring sustained release over prolonged periods of time. (b)

Increased loading efficiency and sustained release from well-dispersed particles could be

obtained by coating calcium phosphate particles with clindamycin adsorbed on them with a

layer of polymer, in this case poly-(D,L-lactide-co-glycolide) (PLGA). High-resolution

transmission electron microscopic images of PLGA-coated hydroxyapatite (c) and

hydroxyapatite nanoparticles (encircled by dashed lines) dispersed in a chitosan matrix

(d).

131

Reprinted with permission from Elsevier (Vukomanović M, Škapin S, Jančar B,

Maksin T, Ignjatović N, Uskoković V, Uskoković D. Poly(D,L-lactide-co-glycolide)/

hydroxyapatite core-shell nanospheres. Part 1: A multifunctional system for controlled drug

delivery. Coll Surf B Biointerfaces. 2011;82(2): 404–13).

Uskoković

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FIG. 6.

An approximate degradation half-life (T

1/2

) for pure poly(lactic acid) (PLA), pure

poly(glycolic acid) (PGA), and their copolymers at various weight ratios (a) and a drastic

difference in the kinetics of drug release from polyethylene glycol diacrylate with 2 different

percentages of cross-linking: 0.5 and 1 (b).

251

Reprinted with permission from Elsevier;

Middleton JC, Tripton AJ. Synthetic biodegradable polymers as orthopaedic devices.

Biomaterials. 2000;21:2334–46.

Uskoković Page 58

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 7.

Schematic description of a hypothetical multifunctional composite nanoparticle (a) and the

subdivision of the tissue regeneration process following injury at multiple stages (b). Stages

specific to the bone regeneration process (c) and adverse outcomes of their obviation or

incompletion (d) are shown.

362,363

Reprinted with permission from American Cancer

Society; Ma X, Zhao Y, Liang XJ. Theranostic nanoparticles engineered for clinic and

pharmaceutics. Acc Chem Res 2011;44(10):1114–22, and Elsevier; Mehta M, Schmidt-

Bleek K, Duda GN, Mooney DJ. Biomaterial delivery of morphogens to mimic the natural

healing cascade in bone. Adv Drug Deliv Rev. 2012;64(12):1257–76.

Uskoković

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 8.

An example of composite monodisperse nanoparticles with simultaneous antioxidative,

antibacterial, and osteoinductive properties. These nanoparticles are formed by coating poly-

(D,L-lactide-co-glycolide) (PLGA) around silver (Ag) poly(glycolic acid) (PGA) core-shell

nanoparticles (Nps) with ascorbic acid dispersed therein. The particles reduced the

concentration of superoxide in human umbilical vein endothelial cells, suppressed the

growth of Escherichia coli and methicillin-resistant Staphylococcus aureus, and upregulated

the expression of 2 osteogenic markers: osteocalcin and protocollagen type I.

392,393

Adapted

and reprinted with permission from Springer (Stevanović M, Savanović I, Uskoković V,

Škapin SD, Bračko I, Jovanović U, Uskoković D. A new, simple, green and one-pot four-

component synthesis of bare and poly(α, γ, L-glutamic acid) capped silver nanoparticles.

Coll Polym Sci. 2012;290(3):221–31) and the American Chemical Society (Stevanović M,

Uskoković V, Filipović M, Škapin SD, Uskoković DP. Composite PLGA/AgNpPGA/AscH

nanospheres with combined osteoinductive, antioxidative and antimicrobial activities. ACS

Appl Mat Interfaces. 2013;5(18):9034–42).

Uskoković Page 60

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FIG. 9.

Hypothetic release curves for a drug delivery device under two different measurement

regimens: daily replacements of a comparatively small volume of the solvent (circles) and

usage of a considerably larger volume of solvent with no daily replenishments (triangles).

Uskoković Page 61

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 10.

(a) A schematic description of the uptake of a calcium phosphate nanoparticle (CaPs) by the

cell and its gene transfection with a DNA plasmid attached to the nanoparticle.

448

(b) A

single-plane confocal optical image of fluorescently stained osteoblastic cell nuclei,

cytoskeletal f-actin, and CaP aggregates containing clindamycin following 48 hours of

incubation. Part A is reprinted with permission from Elsevier (Nouri A, Castro R, Santos JL,

Fernandes C, Rodrigues J, Tomás H. Calcium phosphate-mediated gene delivery using

simulated body fluid (SBF). Int J Pharm. 2012;434(1–2):199–208).

Uskoković Page 62

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Author Manuscript Author Manuscript Author Manuscript Author Manuscript

FIG. 11.

(a) A single-plane confocal optical micrograph of a fluorescently stained calcium phosphate

nanoparticle conglomerate loaded with clindamycin and osteoblastic MC3T3-E1 cells (f-

actin; nucleus) following 21 days of incubation in differentiation medium. This image shows

intimate contact between the cells and the material, a direct indication of the

osteoconductivity of the latter. (b) Results of a gene expression study performed using

quantitative reverse transcriptase polymerase chain reaction and showing diminished

expression of three different osteogenic markers BGLAP (left), Col I (middle), and Runx2

(right) in osteoblastic MC3T3-E1 cells incubated with an antibiotic, clindamycin phosphate

(CL). The effect was compensated for when incubation was carried out in the presence of

either hydroxyapatite nanoparticles (HAP) per se or HAP loaded with CL (HAP/CL).

Messenger RNA expression was detected relative to the housekeeping gene ACTB. *Genes

are significantly upregulated (P < 0.05) with respect to the control group. +Genes are

significantly down-regulated (P < 0.05) with respect to the control group.

453

Reprinted with

permission from John Wiley and Sons (Uskoković V, Desai TA. Phase composition control

of calcium phosphate nanoparticles for tunable drug delivery kinetics and treatment of

osteomyelitis. Part 2: antibacterial and osteoblastic response. J Biomed Mat Res Part A.

2013;101:1427–36).

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Uskoković Page 64

Table 1

Main Calcium Phosphate Phases

103–108

Phase Chemical Formula Space Group

at 37°C

Solubility (mg/dm

)

MCPA Ca(H

)

Triclinic P

1.14

17 · 10

MCPM Ca(H

)

·H

Triclinic P

1.14

18 · 10

DCPD CaHPO

·2H

O Monoclinic I

6.6 88

DCPA CaHPO

Triclinic P

7.0 48

β-CPP Ca

Tetragonal P4

18.5 7.6

ACP Ca

(PO

)

·nH

O / 25 0.8

α-TCP Ca

(PO

)

Monoclinic P2

/a 25.5 2.5

β-TCP Ca

(PO

)

Rhombohedral R3cH 29.5 0.5

TTCP Ca

(PO

)

O Monoclinic P2

37.5 0.7

OA Ca

(PO

)

O Pseudo-hexagonal P6

/m 69 87

CDHA Ca

10−x

(HPO

)

(PO

)

6−x

(OH)

2−x

(0 < x < 1) Pseudo-hexagonal P6

/m 85 9.4

OCP Ca

(PO

)

·5H

Triclinic P

97.4 8.1

HA Ca

(PO

)

(OH)

Pseudo-hexagonal P6

/m 117.3 0.3

FA Ca

(PO

)

Pseudo-hexagonal P6

/m 120 0.2

ACP, amorphous calcium phosphate (data pertain to the phase obtainable at pH 9–11); CDHA, calcium-deficient hydroxyapatite; CPP, calcium

pyrophosphate; DCPA, dicalcium phosphate anhydrous, also known as monetite; DCPD, dicalcium phosphate dihydrate, also known as brushite;

FA, fluoroapatite; HA, hydroxyapatite; MCPA, monocalcium phosphate anhydrous; MCPM, monocalcium phosphate monohydrate; OA,

oxyapatite; OCP, octacalcium phosphate; TCP, tricalcium phosphate; TTCP, tetracalcium phosphate.

Crit Rev Ther Drug Carrier Syst. Author manuscript; available in PMC 2015 April 22.