Interdisciplinary Reviews - Nanomedicine and Nanobiotechnology [Vol.1 Issue 3]

Citation preview

Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology Copyright © 2009 John Wiley & Sons, Inc.

TABLE OF CONTENTS Volume 1 Issue 3 (May/June 2009)

Overviews

Nanomedicine for respiratory diseases (p 255-263)Hulda Swai, Boitumelo Semete, Lonji Kalombo, Paul Chelule, Kevin Kisich, Bob SieversPublished Online: Mar 11 2009 9:56AM DOI: 10.1002/wnan.33

Opinions

Nanoparticle therapeutics: a personal perspective (p 264-271)Scott E. McNeilPublished Online: Jan 13 2009 3:02PM DOI: 10.1002/wnan.6

Advanced Reviews

In vivo visualization of macrophage infiltration and activity in inflammation using magnetic resonance imaging (p 272-298)Nicolau Beckmann, Catherine Cannet, Anna Louise Babin, François-Xavier Blé, Stefan Zurbruegg, Rainer Kneuer, Vincent DoussetPublished Online: Jan 13 2009 11:28AM DOI: 10.1002/wnan.16

Magnetic resonance relaxation properties of superparamagnetic particles (p 299-310)Yves Gossuin, Pierre Gillis, Aline Hocq, Quoc L Vuong, Alain RochPublished Online: Mar 11 2009 10:02AM DOI: 10.1002/wnan.36

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis (p 311-323) Shelton D. Caruthers, Tillmann Cyrus, Patrick M. Winter, Samuel A. Wickline, Gregory M. LanzaPublished Online: Jan 13 2009 3:05PM DOI: 10.1002/wnan.9

Cell-targeting and cell-penetrating peptides for delivery of therapeutic and imaging agents (p 324-335) Rudolph L Juliano, Rowshon Alam, Vidula Dixit, Hyun Min KangPublished Online: Jan 12 2009 3:52PM DOI: 10.1002/wnan.4

Nanotechnology for bone materials (p 336-351)Nhiem Tran, Thomas J. WebsterPublished Online: Mar 11 2009 10:06AM DOI: 10.1002/wnan.23

Overview

Nanomedicine for respiratory diseases Hulda Swai1∗ , Boitumelo Semete1 , Lonji Kalombo1 , Paul Chelule1 , Kevin Kisich2 and Bob Sievers3 Treatment of respiratory diseases and infections has proved to be a challenging task, with the incidence of these ailments increasing worldwide. Nanotechnologybased drug and gene delivery systems offer a possible solution to some of the shortfalls of the current treatment regimen. Nanobased drug delivery systems have revolutionised the field of pharmacotherapy by presenting the ability to alter the pharmacokinetics of the conventional drugs to extend the drug retention time, reduce the toxicity and increase the half-life of the drugs. Delivery of exogenous genes to the airway epithelium in vivo has been limited by several physiological barriers, resulting in the low success rate of these systems. With the advent of nanotechnology, DNA compacted with cationic polymers to produce nanoparticles has exhibited a significant increase in the transfection efficiencies. With nanoparticulate drug/gene delivery systems, specific cells can be targeted by functionalising the polymeric nanoparticles with ligands that allow the particles to dock at a specific site of the cell. In addition, polymeric systems allow for the cargo to be released in a controlled and stimuli-responsive manner. The advantages that nanoparticulate delivery systems present in the treatment of respiratory diseases and infections are summarised in this review .  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 255–263

T

he incidence of respiratory diseases and infections is increasing worldwide. In general, respiratory diseases are physiologically classified as obstructive or restrictive. Obstructive diseases generally impede the flow rate into and out of the lungs, whereas the restrictive conditions cause reduction in the functional volume of the lungs.1 The most common obstructive diseases are asthma, chronic obstructive pulmonary disease (COPD), respiratory allergies, occupational lung diseases and pulmonary hypertension. Currently, 300 million people have asthma, 80 million people have moderate-to-severe COPD while millions of others have mild COPD, allergic rhinitis and other often underdiagnosed chronic respiratory diseases.2,3 Respiratory tract infections, which can be either lower or upper tract infections, are caused by either viral or

∗ Correspondence

to: [email protected]

1 Council

for Scientific and Industrial Research Polymers & Bioceramics, Pretoria, 0001, South Africa 2 National

Jewish Medical and Research Center, Denver, CO, USA

3 AKIN-DRYLLC,

Boulder, CO, USA

DOI: 10.1002/wnan.033

Vo lu me 1, May /Ju n e 2009

bacterial infections. The World Health Organisation (WHO) has classified chronic respiratory diseases as one of the main diseases afflicting the human race and as a result efforts have been dedicated to their prevention, diagnosis and treatment.2 Poor therapeutic outcomes have been reported for a number of treatment regimens for both respiratory diseases and infections, including asthma and tuberculosis (TB), respectively. These can be attributed mainly to patient non-compliance to the prescribed medication, which in many cases is a result of inadequate modes of drug administration.4 The management of the respiratory infections is quite delicate as the early detection of the aetiological agents of the disease appears very challenging.5 A rise in drug-resistant strains of infectious organisms such as Streptococcus pneumoniae and Mycobacterim tuberculosis has created challenges in the treatment of these infections.6–8 S. pneumoniae is known to be resistant to most of the antibiotics used to treat pneumonia, including penicillin, macrolides, doxycycline and trimethoprim–sulfamethoxazole, and resistance to second- and third-generation cephalosporins has also increased.6 Multi-drug-resistant M. tuberculosis

 2009 Jo h n Wiley & So n s, In c.

255

Overview

www.wiley.com/wires/nanomed

is resistant to two of the first-line anti-TB drugs that is, isoniazid and rifampicin.8 Recently, cases of extremely drug-resistant TB (XDR-TB) have been reported in South Africa and other parts of the world, with high prevalence in HIV-positive individuals. XDR-TB is resistant to both first- and second-line anti-TB drugs.9 Other factors that play a role in increasing the rates of infection include allergy and toxic exposures, including tobacco smoke and ambient pollution.3 Treatment of respiratory diseases and infections still poses a great challenge. Much progress has been done with regard to varying the mode of delivery of therapeutic drugs for respiratory diseases and infections. Currently, interest in the use of nanotechnology-based delivery systems has gained momentum. Nanotechnology offers a broad range of opportunities for improving the diagnosis and therapy for respiratory diseases and infections, in particular nanotechnology-based drug delivery systems, which represents an area of particular promise for the treatment of lung diseases. This review presents a summary of chitosan-based nanodrug delivery systems for the treatment of asthma and respiratory infections as well as DNA nanoparticulate delivery systems for the treatment of cystic fibrosis (CF).

CURRENT TREATMENT FOR RESPIRATORY DISEASES To date, the most effective treatment for respiratory diseases resulting in airway inflammation has been oral or injectable corticosteroids administered generally to treat asthma and COPD. Many systemic side effects can occur as a result of the chronic use of corticosteroids; however, much advances have been made in this area, in that corticosteroids can be given by inhalation.9 This route of delivery has minimised systemic absorption of the drugs and many complications previously observed with injectable and oral dosages.10 Although inhalation delivery of the drug has addressed these factors, the persistent challenge is that the lung is functionally and anatomically heterogeneous, thus the dose and drug distribution in the lungs play an important role in reproducible delivery and thus successful therapy.4 Viral respiratory infections such as influenza have no effective and safe antiviral compound and are not susceptible to antibiotic treatment; however antibiotics are generally prescribed for secondary infections. Ribavirin, an antiviral compound with activity against a number of DNA and RNA viruses, has been used to treat viral respiratory infections such as influenza and respiratory syncytial virus (RSV) infection.11 At present oral ribavirin is used in Mexico 256

against influenza, and the aerosol dosage form has been used to treat RSV-related diseases in children. The challenges with ribavirin are the associated adverse side effects, such as haemolytic anaemia, which have been reported to be dose dependent. An additional concern is that this compound has been identified as a tarotogen in some animal species.12 Bacterial respiratory infections on the other hand are treated with oral or injectable antibiotics. Although drugs against respiratory infections such as S. pneumoniae and M. tuberculosis are effective, these drugs generally have to be administered as combination therapy in high doses for long durations of treatment to maintain therapeutic levels and they also have poor bioavailability.6,13 Because of the high doses administered and the associated side effects, patient non-compliance as mentioned above has led to the inefficacy of the treatment regimen. The incorrect dosing of chemotherapy has also been reported to be linked to the emergence of multi-drug-resistant strains,14 and these challenges have posed a need to develop novel ways of delivering the therapeutic compounds.15

ADVANTAGES OF NANOPARTICULATE DRUG DELIVERY SYSTEMS Nanoparticulate drug delivery, also referred to as nanomedicine,16 involves the use of colloidal carrier systems for encapsulation or conjugation of therapeutic compounds to polymeric material. Because of the size of the nanoparticles, which could be in the form of either a nanocapsule or a nanosphere, these particles are able to reach the ‘hard’ to targeted sites of the body. The ability to target the lung via inhalable micro- or nanoparticles has the potential of minimising drug resistance, reducing side effects and also lowering the therapeutic dose which is usually administered orally.17 The entrapment of the drugs in polymeric particles allows drug release from the polymeric material in a predesigned manner which may be either constant over a long period, cyclic, or it may also be triggered by the cellular microenvironment or other external events, such as pH, temperature, oxidative conditions or an external magnetic field.18 Functional groups on the surface of the particles allow chemical conjugation of various ligands, peptide, DNA and sugars to enable active targeted delivery of particles to cells where these molecules will be recognised.19 In addition, in order to make the particles stealth, polymers such as polyethylene glycol (PEG) and polyvinypyrrolidone (PVP) have been attached onto

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanomedicine for respiratory diseases

their surface. These polymers have been reported to minimise opsonisation, thus increasing the residence time of the particles in the blood circulation.20 In recent years, biodegradable polymers such as poly(d,l-lactide), poly(lactic acid) (PLA), poly(d,lglycolide) (PG), poly(lactide-co-glycolide) (PLG) and poly(cyanoacrylate) (PCA), as well as chitison, alginate and gelatine to name a few that have attracted considerable attention as potential drug delivery devices in view of their applications in the control release of drugs, their ability to target particular organs/tissues as carriers of DNA in gene therapy and in the ability to deliver proteins, peptides and genes.21

CHITOSAN NANOSPHERES FOR TREATMENT OF ALLERGIC ASTHMA Asthma is a chronic disease characterised by allergeninduced airway inflammation resulting in the infiltration of inflammatory cells such as eosinophils and epithelial hyperplasia leading to hypersecretion of mucus and the presence of airway hyperresponsiveness (AHR) to a variety of environmental stimuli.22–24 The abnormal low airflow rates in the airways can be partially or fully restored by prescription of bronchodilator and anti-inflammatory medications.25 Traditional treatments for bronchial asthmatic inflammation include the administration of non-steroidal anti-inflammatory drugs, whereas for more advanced stages of disease, immunosuppressive agents such as methotrexate, cyclosporin and azathioprine have been extensively used.26 In addition, theophylline has been used for a long time as a frontline drug for the treatment of asthma, but its use has presented some side effects.22,24 Despite these drawbacks, theophylline remains the widely prescribed anti-asthmatic agent.27–29 In line to improve the efficacy of the suggested therapy, some new active agents have been identified and have shown some benefits. They consist mainly of gene-based drugs. In this category, interferon-gamma (IFN-γ ) has emerged as an excellent candidate for asthma therapy.23,30–32 Although the administration of recombinant IFN-γ showed its efficacy in tackling the airway disorders and inflammations in murine models,33,34 its short half-life in vivo as observed with other cytokine-based drugs requires repeated and frequent dosing which can result in anti-drug antibodies that block therapeutic effectiveness over time.34 In order to circumvent these shortfalls, Kumar and co-workers investigated systematically the possibility of using IFN-γ gene transfer which inhibits both antigenand Th2-induced pulmonary eosinophilia and airway Vo lu me 1, May /Ju n e 2009

hyper-reactivity.23 For this purpose, a plasmid DNA (pDNA)-encoded IFN-γ gene transfer was developed. Because of the challenges that gene delivery poses, such as inefficient transfection efficiencies, the use of appropriate delivery systems to ensure that the DNA reaches the target site becomes imperative. Hence, cationic polymers have been proposed as a promising approach to the development of non-viral vectors.32 Chitosan, a cationic natural biopolymer produced by the alkaline N-deacetylation of chitin, has become an interesting material in pharmaceutical application, especially as an inhalation drugcarrier,35,36 due to its biodegradability,37 biocompatibility, low toxicity as well as the absence of immunogenic issues.38,39 It has shown the ability to form stable complexes, also known as polyplexes, with DNA via electrostatic interactions 39,40 and efficiently deliver the gene intracellularly.40,41 The efficiency of the intracellular gene expression of such systems has been attributed mainly to the inherent properties of chitosan including its mucoadhesivity and its polycationic nature.40

CHITOSAN NANOPARTICLES AS DELIVERY VEHICLES Chemically modified or unmodified chitosan nanoparticles (CINs) have been intensively investigated as they present enhanced potential for gene delivery.35 Among them, thiolated chitosan nanoparticles (TCNs) have recently been short-listed as excellent candidates for drug delivery to the mucus-rich bronchial epithelium.23,42,43 Lee et al.28 demonstrated the effectiveness of thiolated chitosan as delivery system for theophylline. When administered to ovalbuminchallenged BALB/c mice, it reduced the number of eosinophils in bronchoalveolar lavage (BAL) fluid.28 These results were in agreement with the findings of several investigators in the same field.43–45 On the other hand, Kumar et al.42 have shown that a pDNAencoded IFN-γ complexed with CINs, intranasally administered into a BALB/c mouse model of allergic asthma, has been able to reduce the hypersensitivity, suggesting that a high efficiency of gene transfection in the lung epithelial cells 23,46 has been achieved due to enhanced mucoadhesiveness of TCNs.

CHITOSAN NANOSPHERES FOR PROTECTION AGAINST RESPIRATORY INFECTION Very recently, a special interest has been directed to the development of an RNA interference (RNAi)

 2009 Jo h n Wiley & So n s, In c.

257

Overview

www.wiley.com/wires/nanomed

(a)

Magnet

Magnetic nanoparticles

(b)

Surface of chest Magnetic nanoparticles

Fmag

Magnet

Magnetic flux lines

Mucus layer Lung epithelium

FIGURE 1 | Schematic illustration of targeted delivery of genes to the lung via magnetic nanoparticle based delivery. (a) Illustrates attraction of the inhaled magnetic nanoparticles to the high gradient magnets. (b) A section view illustrating magnetic particles in contact with the mucus lining of the lung. (Reprinted with permission from Ref 70. Copyright 2006 Keele University).

targeted to a target transcript that encodes a protein involved in development, pathogenesis or symptoms of an IgE-mediated disease or condition such as allergic rhinitis or asthma.47 The mechanism of RNAi generally involves cleavage of the target RNA or inhibition of its translation product. The challenge, with RNAi therapeutic systems, has been primarily the delivery of the RNA into the cytosol. Because of the aforementioned mucoadhesive properties of chitosan, and the ability to deliver the interfering RNA into the cell when RNA is complexed with chitosan into a nanoparticle, these gene delivery systems have been successfully applied to oral and nasal route gene therapy systems extending beyond asthma treatment.48–50 Respiratory syncytial virus has been recognised as the leading cause of severe bronchiolitis and pneumonia in infants worldwide and also as resulting in lower respiratory tract infections in immunodeficient and elderly adults.48,49 Natural immunity 258

to RSV is incomplete, and infection recurs throughout life.49 Kong et al.48 in their report on the prophylactic effects of short interfering non-structural proteins (siNS1) construct in preventing RSV infection in rats, have illustrated that the siNS1 treatment reduced RSV titres significantly and prevented the accompanying lung damage and airway hyper-reactivity when pDNAs expressing anti-NS1 RNA or an unrelated sequence were complexed with CINs and instilled intranasally 1 day prior to intranasal infection with RSV.48 Similar results in a murine model confirmed an effective prophylactic effect of a mucosal gene expression vaccine (GXV) made up of a cocktail of at least four different pDNAs encoding corresponding RSV antigens, coacervated with chitosan to formulate nanospheres.50,51 The intranasal administration with GXV resulted in significant induction of RSVspecific antibodies, nasal IgA antibodies, cytotoxic T lymphocytes and IFN-γ production in the lung and splenocytes.51 A similar effect in the reduction of allergic response was reported in ovalbulim-sensitised mice that were administered chitosan/pIFN-γ nanoparticles prior to an asthma-inducing challenge.41 Subsequent to prophylactic treatment with chitosan/pIFN-γ particles, splenocytes collected from treated mice exhibited increased secretion of IFN-γ and decreased secretion of IL-5 and IL-4.41 Amidi et al.52 found that monovalent influenza subunit vaccine-loaded N-trimethyl chitosan (TMC) nanoparticles were an effective carrier system for nasal delivery. Furthermore, they observed that the immune responses elicited by the antigenloaded TMC nanoparticles were likely attributed to cellular uptake of the nanoparticles in the nasal epithelium and nasal-associated lymphoid tissues (NALT) and subsequent access of the vaccine to sub-mucosal lymphoid tissues.52 Bivas-Benita et al.53 used an HLA-A2 transgenic mouse model to investigate the effects of pulmonary delivery of a new DNA plasmid encoding eight HLA-A*0201-restricted T-cell epitopes from M. tuberculosis formulated in CINs. They have shown that pulmonary administration of the DNA plasmid incorporated in CINs induced increased levels of IFNγ secretion. Maturation of dendritic cells was also observed when compared with pulmonary delivery of plasmid in solution and the more frequently used intramuscular immunisation route,53 in line with observations of other groups.54,55 The studies summarised above indicate the advancement of research in the area on nanoparticulate gene delivery systems. The success of polymer compacted DNA in the treatment of asthma has also been reported for CF, a disease where gene therapy has been explored over a decade ago.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanomedicine for respiratory diseases

NANOTECHNOLOGY FOR TREATMENT OF CYSTIC FIBROSIS Cystic fibrosis is a hereditary autosomal recessive disease of the lungs that also affects the digestive system, resulting in progressive disability and early death. Thick mucous production in the airways leads to distorted muco-ciliary clearance, a less competent immune system and inflammation, resulting in frequent lung infections,56 such as allergic bronchopulmonary aspergillosis and mycobacterium avium complex. CF is caused by mutations in the CF transmembrane conductance regulator (CFTR) gene. The product of the gene is a chloride ion channel, which is important for controlling the movement of the chloride ions from outside the cell into the cytoplasm. When the protein is not functioning, the chloride ions are trapped outside the cell. Protons such as sodium interact with the chloride ions producing sodium chloride salt. It is hypothesised that the lack of chloride transport leads to the accumulation of nutrient-rich mucus in the lungs, which allows propagation of bacteria in these sites.57 The most common mutation in the CFTR gene is a deletion in position 508 of the amino acid sequence leading to the deletion of a phenylalanine in the protein. This mutation accounts for 70% of worldwide CF cases. More than 500 mutations have been associated with CF.58 CF is currently treated with antibiotics that are prescribed on the basis of the infectious agent; however, many of these bacteria are resistant to multiple antibiotics and require weeks of treatment with intravenous antibiotics such as tobramycin, ciprofloxacin and piperacillin. Inhaled therapy with other antibiotics is also followed in some cases to improve lung function by impeding the growth of colonised bacteria.59,60 Oral therapy is also sometimes administered to prevent infection or to control the current infection. In addition to the prolonged treatment time, these antibiotics manifest with side effects such as hearing loss and kidney failure. To address these shortfalls, Sweeney

et al.61 have reported a liposomal formulation of ciprofloxacin powder manufactured using a sprayfreeze drying process with the required mass mean aerodynamic diameter and fine particle fraction. This system is postulated to be efficient as inhaler and thus increases the bioavailability of the drug.61 A new form of therapy that is gaining much momentum is gene therapy. Efforts in this area were previously centred on the transfection of airway epithelial cells with the normal CFTR gene using the adenovirus vector, adeno-associated virus (AAV) and lentivirus. Adenovirus, although natural targets of the airway epithelium, led to unanticipated host immune response to the vector,62 which led to the termination of the human trial. An additional shortfall was transient gene expression. The AAV vectors, although safe, require a helper virus to replicate; however, they are capable of site-directed insertion into DNA, reducing the risk of insertional mutagenesis.63 Lentiviruses have the ability of transfecting cells that are not terminally differentiated such as the basal or airway progenitor cells, which make them a good candidate for CF gene therapy; however, the shortfall is still insertional mutagenesis which causes a significant risk.64 To date, much effort is put into synthetic gene delivery systems that address the factors that were not completely achieved with the aforementioned systems. These include non-immunogenecity of the vector, stability during processing and safety when the cells are transfected; thus the vector should stably integrate into the progenitor cell genome or be safe and effective with repeated administration. Another barrier that led to the inefficiency of these vectors is the inefficient deposition of the vector on the apical surface of the airway epithelium due to the thick and viscous mucus that presents a physical barrier.65 Viral systems are by far the most efficient in DNA delivery; however, great advances have been made with cationic liposomal systems,66 DNA-binding peptides and polymeric systems,67 which incorporate

FIGURE 2 | (a) Compacted DNA containing the (a)

Vo lu me 1, May /Ju n e 2009

(b)

 2009 Jo h n Wiley & So n s, In c.

luciferase gene. (b) Bioluminescent image of mice injected with the DNA. (Personal Communication: A. Ziady, Ph.D. at Case Western Reserve University).

259

Overview

www.wiley.com/wires/nanomed

various biological parameters that enable receptormediated endocytosis, promote endosomal escape after entry into the cell and lead to a cytoplasmic selfreplication of the gene. The next section will review the success made with these systems and future trends.

LIPOSOMAL SYSTEMS AND DNA NANOSPHERES Liposomal systems where the drug is encapsulated into the phospholipid sphere have previously been utilised.61 Because of the lipid composition of the spheres, these liposomal formulations can readily pass through the cellular membrane; however, the drawback with these systems is that these spheres are generally too large (>100 nm) to pass through the nuclear membrane pores. Nuclear incorporation can occur in actively dividing cells, during cell division, but only for slow dividing cells such as the lung basal epithelium.61 Such limitation has led to the investigation of alternative nanoparticulate gene delivery systems. Truong-Le’s group at Johns Hopkins has prepared DNA cross-linked with gelatine via electrostatic interaction between the two components to form nanosphere coacervates.67 These yielded a high DNA encapsulation efficiency and loading when sodium sulphate was used as a desolving agent. To increase their cellular uptake, transferrin was attached to the surface of the nanospheres, and the encapsulation of chloroquine led to further transfection efficiency.67 Chloroquine was reported to play a role in the acidification of the endosomes, thus improving the cellular delivery of the DNA and contributing towards the stability of the DNA.67 When cells defective of a chloride transport channel were treated with CFTR DNA nanospheres, successful complementation of the channel was achieved, similar to mutant cells treated with AAV.67 Such promising results with DNA nanoparticles were also obtained by Ziady et al.,68 where they also illustrated efficient DNA uptake and expression with CFTR DNA compacted with PEG–lysine to form nanocomplexes of sizes ranging from 20 to 25 nm.68,69 Results from animal studies with this system showed evidence of partial-to-complete correction of the defective CFTR gene.68 The group also attached a 17 amino acid ligand for a serine

proteinase inhibitor (serpin)-enzyme complex receptor to these nanospheres. This modification facilitated efficient targeted entry of the particles.68,69

MAGNETIC NANOPARTICULATE DELIVERY SYSTEMS Recent work by the UK Cystic Fibrosis Gene Therapy Consortium has focussed on magnetic particles with either the therapeutic gene or a reported gene attached to them. These particles are inhaled and targeted to the airway epithelium via positioning of a strong, high gradient magnet over the target side as indicated in Figure 1, which functions to pull the particles into contact with the cells.70 To improve the in vivo transfection efficiency of DNA delivery of this system, the group is currently developing an oscillating magnet array system, which will introduce energy and a lateral component to advance the movement and interaction of the particles with the epithelial cells.71

CONCLUSION One of the challenges facing gene therapy for respiratory diseases and infections is determining the dose frequency. Technologies where nanoparticles can be visualised and gene expression be measured once the nanoparticles are administered have recently been developed.72 As indicated in Figure 2, the particles with a reporter gene, in this case the firefly luciferase, can be followed via bioluminescent imaging. This approach will enable in vivo tracking of the DNA complexes and determination of when the next dose is to be administered. Further challenges will however be determining a system that is sensitive enough to measure a range of gene expression levels, developing a tag to enable in vivo imaging and also targeting the nanoparticles to specific/diseased cell types.69,73 Although nanoparticulate gene therapy and drug delivery hold much promise, the toxic effects of nanoparticles still need to be thoroughly investigated. Various properties of the nanoparticulate systems such as chemistry, size and physical properties will have to be assessed as they may play a role in the toxic effects of nanoparticles.

REFERENCES 1. Respiratory tract infections: www.healthinsite.gov.au/ topics/Respiratory Tract Diseases (accessed January 2009).

260

2. World Health Organization. WHO strategy for prevention and control of chronic respiratory diseases. Geneva: World Health Organization; 2002.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanomedicine for respiratory diseases

3. Enarson DA, Chretien J. Epidemiology of respiratory infectious diseases. Curr Opin Pulm Med 1999, 5(3):128.

18. Schmoljohann D. Thermo- and pH- responsive polymers for drug delivery. Adv Drug Deliv Rev 2006, 58:1655–1670.

4. Gonda I, Schuster JA, Rubsamen RM, Lloyd P, Cipolla D, et al. Inhalation delivery systems with compliance and disease management capabilities. J Control Release 1998, 53:269–274.

19. Torchilin V. Multifunctional nanocarriers. Adv Drug Deliv Rev 2006, 58:1523–1555.

5. Ruiz M, Ewig S, Marcos MA, Martinez JA, Arancibia F, et al. Etiology of community-acquired pneumonia: impact of age, co-morbidity, and severity. Am J Respir Crit Care Med 1999, 160:397–405. 6. Butler JC, Hofmann J, Cetron MS, Elliott JA, Facklam RR, et al. The continued emergence of drugresistant Streptococcus pneumoniae in the United States: an update from the Centers for Disease Control and Prevention’s Pneumococcal Sentinel Surveillance System. J Infect Dis 1996, 174:986–993. 7. Kunin CM. Resistance to antimicrobial drugs–a worldwide calamity. Ann Intern Med 1993, 118:557–561.

20. Evora C, Soriano I, Rogers RA, Shakesheff KM, Hanes J, et al. Relating the phagocytosis of microparticles by alveolar macrophages to surface chemistry: the effect of 1,2-dipalmitoylphosphatidylcholine. J Control Release 1998, 51:143–152. 21. Kumaresh S, Soppimath M, Aminabhavi A, Kulkarni R, Langer R. Biomaterials in drug delivery and tissue engineering: one laboratory’s experience. Acc Chem Res 2000, 33:94–101. 22. Lim S, Tomita K, Carramori G, Jatakanon A, Oliver B, et al. Low-dose theophylline reduces eosinophilic inflammation but not exhaled nitric oxide in mild asthma. Am J Respir Crit Care Med 2001, 164:273–276.

8. Petrini B, Hoffner S. Drug resistant and multi-drug resistant tubercle bacilli. Int J Antimicrob Agents 1999, 13:93–97.

23. Kumar M, Kong X, Behera AK, Hellermann GR, Lockey RF, et al. Chitosan IFN-r-pDNA nanoparticle (CIN) therapy for Allergic Asthma. Genet Vaccines Ther 2003, 1:3.

9. Wright A. Emergence of Mycobacteria Tuberculosis with extensive resistance to second line drugs worldwide. JAMA 2006, 295:2349–2351.

24. Barnes PJ. Theophylline - new perspectives for an old drug. Am J Respir Crit Care Med 2003, 167:813–818.

10. Todd GR, Acerini CL, Ross-Russell R, Zahra S, Warner JT, et al. Survey of adrenal crisis associated with inhaled corticosteroids in the United Kingdom. Arch Dis Child 2002, 87:457–461. 11. Smith RA, Kirkpatrick W. Ribavirin: Structure and Antiviral Activity Relationships Ribavirin: Abroad Spectrum Antiviral Agent. New York, NY: Academic Press; 1980, 1–21. 12. Bani-Sadr F, Carrat F, Pol S. Risk factors for symptomatic mitochondrial toxicity in HIV/hepatitis C viruscoinfected patients during interferon plus ribavirinbased therapy. J Acquir Immune Defic Syndr 2005, 40:47–52. 13. du Toit LC, Pillay V, Danckwerts MP. Tuberculosis chemotherapy: Current drug delivery approaches. Respir Res 2006, 7:118. 14. Lipsitch M, Samore MH. Antimicrobial use and antimicrobial resistance: a population perspective. Emerg Infect Dis 2002, 8:347–354. 15. Prabakaran D, Singh P, Jaganathan KS, Vyas SP. Osmotically regulated asymmetric capsular systems for simultaneous sustained delivery of anti-tubercular drugs. J Control Release 2004, 95(2):239–248. 16. Duncan R. Nanomedicine gets clinical. NanoToday 2005, 8:16–17. 17. Pandey R, Khuller GK. Solid lipid particle based inhalable sustained drug delivery against experimental tuberculosis. Tuberculosis 2005, 85:227–234.

Vo lu me 1, May /Ju n e 2009

25. Smolensky MH, Lemmer B, Reinberg AE. Chronobiology and chronotherapy of allergic rhinitis and bronchial asthma. Adv Drug Deliv Rev 2007, 59(9–10):823–824. 26. Hedley ML. Gene therapy of chronic inflammatory disease. Adv Drug Deliv Rev 2000, 44:195–207. 27. Sullivan P, Bekir S, Jaffar Z, Page C, Jeffery P, et al. Anti-inflammatory effects of low-dose oral theophylline in atonic asthma. Lancet 1994, 343:1006–1008. 28. Lee D-W, Shirley SA, Lockey RF, Mohapatra SS. Thiolated chitosan nanoparticles enhance anti-inflammatory effects of intranasally delivered theophylline. Respir Res 2006, 7(1):112. 29. Asada M, Takahashi H, Okamoto H, Tanino H, Danjo K. Theophylline particle design using chitosan by the spray drying. Int J Pharm 2004, 270:167–174. 30. Ford JG, Rennick D, Donaldson DD, Venkayya R, McArthur C, et al. IL-13 and IFN-gamma: interactions in lung inflammation. J Immunol 2001, 167:1769–1777. 31. Daines MO, Hershey KGK. A novel mechanism by which interferon-gamma can regulate IL-13 responses: evidence for intracellular stores of IL-13 receptor alpha 2 and their rapid mobilization by interferon-gamma. J Biol Chem 2002, 277(12):10387–10393. 32. Pierkes M, Bellinghausen I, Hultsch T, Metz G, Knop J, et al. Decreased release of histamine and sulfidoleukotrienes by human peripheral blood leukocytes after wasp venom immunotherapy is partially due to induction of IL-10 and IFN-gamma production of T cells. J Allergy Clin Immunol 1999, 103:326–332.

 2009 Jo h n Wiley & So n s, In c.

261

Overview

www.wiley.com/wires/nanomed

33. Flaishon L, Topilski I, Shoseyov D, Hershkoviz R, Fireman E, et al. Cutting edge: anti-inflammatory properties of low levels of IFN-gamma. J Immunol. 2002, 168:3707–3711. 34. Yoshida M, Leigh R, Matsumoto K, Wattie J, Ellis R, et al. Effect of interferon-gamma on allergic airway responses in interferon-gamma-deficient mice. Am J Respir Crit Care Med 2002, 166:451–456. 35. Kim T-H, Jiang H-L, Jere D, Park I-K, Cho M-H, et al. Chemical modification of chitosan as a gene carrier in vitro and in vivo. Prog Polym Sci 2007, 32(7):726–753. 36. Okamoto H, Nishida S, Todo H, Sakakura Y, Lida K, et al. Pulmonary gene delivery by chitosan-pDNA complex powder prepared by a supercritical carbon dioxide process. J Pharm Sci 2003, 92:371–380. 37. Williams RO, Barron MK, Alonso MJ, RemunanLopez C, III Investigation of a pMDI system containing chitosan microspheres and P134a. Int J Pharm 1998, 174:209–222. 38. Lueßen HL, Rentel C-O, Kotze AF, Lehr C-M, de Boer AG, et al. Mucoadhesive polymers in peroral peptide drug delivery. IV. Polycarbophil and chitosan are potent enhancers of peptide transport across intestinal mucosa in vitro. Int J Pharm 1997, 45:15–23. 39. Lee KY, Kwon IC, Kim YH, Jo WH, Jeong SY. Preparation of chitosan self-aggregates as a gene delivery system. J Control Release 1998, 51:213–220. 40. Mohapatra, SS. Chitosan-microparticles for ifn gene delivery, US Patent 20070116767. 41. Dang JM, Leong KW. Natural polymers for gene delivery and tissue engineering. Adv Drug Deliv Rev 2006, 58:487–499. 42. Kumar M, Behera AK, Lockey RF, Zhang J, Bhullar G, et al. Intranasal gene transfer by chitosan-DNA nanospheres protects BALB/c mice against acute respiratory syncytial virus infection. Hum Gene Ther 2002, 13:1415–1425. 43. Kumar M, Behera AK, Lockey RF, Vesely DL, Mohapatra SS. A trial natriuretic peptide gene transfer by means of intranasal administration attenuates airway reactivity in a mouse model of allergic sensitization. J Allergy Clin Immunol 2002, 110:879–882. 44. Caramori G, Adcock I. Pharmacology of airway inflammation in asthma and COPD. Pulm Pharmacol Ther 2003, 16:247–277. 45. Kobayashi M, Nasuhara Y, Betsuyaku T, Shibuya E, Tanino Y, et al. Effect of low-dose theophylline on airway inflammation in COPD. Respirology 2004, 9:249–254. 46. Behera AK, Kumar M, Lockey RF, Mohapatra SS. 2’– 5′ Oligoadenylate synthetase plays a critical role in interferon-gamma inhibition of respiratory syncytial virus infection of human epithelial cells. J Biol Chem 2002, 277:25601–25608.

262

47. Chen, J, Eisen, HN, Ge, Q. RNAi-based therapeutics for allergic rhinitis and asthma, United States Patent 20060058255. 48. Kong X, Zhang W, Lockey RF, Auais A, Piedimonte G, et al. Respiratory syncytial virus infection in Fischer 344 rats is attenuated by short interfering RNA against the RSV-NS1 gene. Genet Vaccines Ther 2007, 5:4. 49. Xie C, He J-S, Zhang M, Xue S-L, Wu Q, et al. Oral respiratory syncytial virus (RSV) DNA vaccine expressing RSV F protein delivered by attenuated Salmonella typhimurium. Hum Gene Ther 2007, 18:746–752.. 50. Shahiwala A, Vyas TK, Amiji MM. Nanocarriers for systemic and mucosal vaccine delivery. Recent Pat Drug Deliv Formul 2007, 1:1–19. 51. Mohapatra, SS, Kumar, M, Huang, S, Leong, K. United States Patents 2003068333. 52. Amidi M, Romeijn SG, Verhoef JC, Junginger HE, Laura BL, et al. N-Trimethyl chitosan (TMC) nanoparticles loaded with influenza subunit antigen for intranasal vaccination: biological properties and immunogenicity in a mouse model. Vaccine 2007, 25:144–153. 53. Bivas-Benita M, van Meijgaarden KE, Franken KLMC, Hans E, Junginger HE, et al. Pulmonary delivery of chitosan-DNA nanoparticles enhances the immunogenicity of a DNA vaccine encoding HLA-A*0201restricted T-cell epitopes of Mycobacterium tuberculosis. Vaccine 2004, 22:1609–1615. 54. Eyles JE, Spiers ID, Williamson ED, Alpar HO. Analysis of local and systemic immunological responses after intra-tracheal, intra-nasal and intra-muscular administration of microsphere co-encapsulated Yersinia pestis sub-unit vaccines. Vaccine 1998, 16(20):2000–2009. 55. Lagranderie M, Ravisse P, Marchal G. BCG-induced protection in guinea pigs vaccinated and challenged via the respiratory route. Tuber Lung Dis 1993, 74(1):38–46. 56. Groneberg DA, Eynot PR, Lim S, Oates T, Wu R, et al. Expression of respiratory mucin in fatal status asthnaticus and mild asthma. Histopathology 2002, 40:367–373. 57. Rosenstein BJ, Zeitlin PL. Cystic fibrosis. Lancet 1998, 351:277–282. 58. Stern RC. The diagnosis of cystic fibrosis. N Engl J Med 1997, 336:487–491. 59. Pai VB, Nahata MC. Efficacy and safety of aerosolised tobromycin in cystic fibrosis. Pediatr Pulmonol 2001, 32:314–327. 60. Westernman EM, Le Brun PP, Touw DJ, Frijlink HW, Heijerman HG. Effect of nebulised colistin sulphate and colistin sulphomethane on lung function in patients with cystic fibrosis: a pilot study. J Cyst Fibros 2004, 3:23–28. 61. Sweeney LG, Wang Z, Loebenberg R, Wong J, Lange CF, et al. Spray-freeze-dried liposomal

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanomedicine for respiratory diseases

ciproflaxacin powder for inhaled aerosol drug delivery. Int J Pharm 2005, 305:180–185. 62. Knowles MR, Hohneker KW, Zhou Z. A controlled study of adenovirus-vector-mediated gene transfer in the nasal epithelium of patients with cystic fibrosis. N Engl J Med 1995, 333:823–832. 63. Moss RB, Rodman D, Spencer LT. Repeated adenoassociated virus serotype 2 aerosol-mediated cystic fibrosis transmembrane regular gene transfer to the lungs of patients with cystic fibrosis: a multi-centre double-blind, placebo-controlled trial. Chest 2004, 125:509–521. 64. Copreni E, Penzo M, Carrabino S, Conense M. Lentivirus-mediated gene transfer to the respiratory epithilium: a promising approach to gene therapy of cystic fibrosis. Gene Ther 2004, 11:S67–S75. 65. Flotte TR, Laube BL. Gene therapy in cystic fibrosis. Chest 2001, 120:124S–131S. 66. Rochat T, Morris MA. Gene therapy for Cystic fibrosis by means of aerosol. J Aerosol Med 2002, 15:229–223. 67. Truong-Le VL, Walsh SM, Schwiebert E, Mao H-Q, Guggino WB, et al. Gene transfer by DNA-gelatin nanospheres. Arch Biochem Biophys 1999, 361:47–56.

68. Ziady AG, Gedeon CR, Muhammad O, Stillwell V, Oette S, et al. Minimal toxicity of stabilised compacted DNA in the murine lung. Mol Ther 2003, 8:948–956. 69. Ziady AG, Gedeon CR, Miller T, Quan W, Payne JM, et al. Transfection of airways epithelium by stable PEGylated poly-lysine DNA nanoparticles in vivo. Mol Ther 2003, 8:936–947. 70. Dobson J. Magnetic nanoparticles-based gene delivery. Gene Ther 2006, 13:283–287. 71. Xenariou S, Griesenbach U, Ferrari S. Using magnetic forces to enhance non-viral gene transfer to airway epithelium in vivo. Gene Ther 2006, 13:1445–1452. 72. Kotlarchyk, M, Lee, Z, Cooper, M, Davis, PB, Ziady, AG. Imaging of sec-R directed and PEG-stabilised gene transfer nanoparticles in CF mice. 8th Annual Meeting of the American Society of gene therapy, 1–5 June. St Louis, MO; 2005. 73. Konstan MW, Davis PB, Wagener JS, Hilliard KA, Stern RC, et al. Compacted DNA nanoparticles administered to the nasal mucosa of cystic fibrosis subjected are safe and demonstrate partial to complete cystic fibrosis transmembrane regulator reconstitution. Hum Gene Ther 2004, 15:1255–1212.

FURTHER READING Freitas RA. NanoMedicine, Vol. 1: Basic Capabilities. Landes Bioscience 1999. Liu F. Huang, L. Development of non-viral vector for systematic gene delivery. J Control Release 2002 78 259–266. Edwards DA, Ben-Jebria A, Langer R.Recent advances in pulmonary drug delivery using large porous inhaled particles. J Appl Physiol 1998 84 379–385. Edwards DA, Dunbar CA.Therapeutic aerosol bioengineering. Annu Rev Biomed Eng 2002 4 93–107. Pandey R, Sharma S, Khuller GK.Oral solid lipid nanoparticle-based antitubercular chemotherapy. Tuberculosis 2005 85 415–420. Du Toit, LC, Pillay V, Danckwerts MP.Tuberculosis Chemotherapy: current drug delivery approaches. Respir Res 2006 7 117–130.

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

263

Opinion

Nanoparticle therapeutics: a personal perspective Scott E. McNeil∗ Nanotechnology offers many potential benefits to cancer research through passive and active targeting, increased solubility/bioavailablility, and novel therapies. However, preclinical characterization of nanoparticles is complicated by the variety of materials, their unique surface properties, reactivity, and the task of tracking the individual components of multicomponent, multifunctional nanoparticle therapeutics in in vivo studies. There are also regulatory considerations and scaleup challenges that must be addressed. Despite these hurdles, cancer research has seen appreciable improvements in efficacy and quite a decrease in the toxicity of chemotherapeutics because of ‘nanotech’ formulations, and several engineered nanoparticle clinical trials are well underway. This article reviews some of the challenges and benefits of nanomedicine for cancer therapeutics and diagnostics.  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 264–271

E

arlier this year I was invited to observe a clinical trial for a nanoparticle formulation of a cancer drug. I run a lab for the National Cancer Institute (NCI) that conducts preclinical characterization of these nanoparticles, and we had been involved with this formulation for a few months. This particular study was resurrecting a drug that had been discontinued in clinical trials 10 years earlier because of immunotoxicity. I am not a physician, so the trial coordinator cleared my presence with the patient several weeks earlier. The patient, I shall refer to her as ‘Sally’ for the purposes of this article, was quite receptive toward allowing a researcher to observe the trial. Sally had been treated with a variety of chemotherapeutics, and had experienced the adverse side effects common to cancer drugs. In the context of cancer treatment, chemotherapy generally refers to the treatment with cytotoxic drugs such as vincristine, methotrexate, doxorubicin, cisplatin, and others. Cytotoxic drugs are effective at killing cancer cells and are the workhorse of most cancer therapy (along with surgery and radiation), but they work by killing neoplastic cells marginally better than they kill other proliferating cells—hopefully killing the cancer without killing the patient. Patients

∗ Correspondence

to: Scott E. McNeil, SAIC-Frederick. E-mail: [email protected]

Nanotechnology Characterization Lab, Imaging and Nanotechnology Group, SAIC-Frederick Inc./National Cancer Institute at Frederick, Frederick, MD, USA DOI: 10.1002/wnan.006

264

treated with cytotoxic drugs commonly suffer from nausea, diarrhea, immunocompromise, neuropathy, and fatigue. Their treatment often has to be interrupted to allow them time to recover from these adverse events. My training in cancer research translates these side effects into acronyms, terms, and numbers that seem sterile and analytic—STD10 , MTD, therapeutic index—but that is simply because ‘controlled poisoning’ is not a term we generally use in scientific textbooks. Sally had volunteered for the clinical trial with the hope of contributing to the development of something more effective at treating cancer, and less apt to cause these pernicious, sometimes lethal, side effects. Her doctors are hoping for the same thing—a solution that targets cancer cells and tumors, with diminished damage to healthy cells. So are oncologists and researchers from academia, government, and industry. So is the NCI. Mortality as a result of cancer is second only to heart disease in the USA. Although novel anticancer drugs such as Gleevec and Herceptin have recently demonstrated a remarkable success, the death rate because of cancer has not decreased in the last 50 years. This year over 1,300,000 Americans will be diagnosed with cancer and over half a million will die from the disease. Buried in those statistics are our relatives, coworkers, friends, and neighbors. Compare those numbers with the progress in heart disease treatment for the same period, where the death rate has dropped by more than half. A report detailing

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanoparticle therapeutics

those bleak figures caught the attention of NCI’s leadership prompting them to launch several new strategic initiatives—involving proteomics, biomarkers, biospecimens, and nanotechnology—to increase our effectiveness at combating cancer.

NANOTECHNOLOGY AS MEDICINE: BEGINNINGS Four years prior to observing Sally’s clinical trial, I had received a call from a contact at NCI-Frederick asking me if I would be interested in helping them structure the Cancer Nanotechnology Plan, and the subsequent Alliance for Nanotechnology in Cancer.1 These programs were initiated to develop basic nanotech research, and then transition the resulting therapeutics and diagnostics into clinical applications. At the time of the call I was understandably skeptical. The notion of injecting ‘multifunctional’, engineered nanoparticles into people to target cancer cells seemed like something out of a 1960s-era SciFi movie. But then I was introduced to the nanotech projects that NCI had sponsored for the past several years, under the Unconventional Innovations Program (UIP). I was given the opportunity to scrutinize the investigators’ claims and vet the technology for myself. I was surprised to find improvements in the safety and efficacy that were not

simply incremental enhancements, but often orders-ofmagnitude improvements compared with their legacy counterparts. Even then, the potential benefits of nanotechnology to cancer therapy were noteworthy, and significant research was being conducted by academics and small biotech firms. Baker’s group at the University of Michigan demonstrated that methotrexate coupled with folic-acid-targeted PAMAM dendritic polymers had improved antitumor activity and decreased toxicity to levels not achievable by the free drug.2 Alnis Biosciences, Inc., have shown that a polymer-coated doxorubicin-nanoparticle formulation delivered high levels of doxorubicin to tumors.3 West and Halas at Rice University were able to target gold/silicon nanoparticles coated with antibodies to breast carcinoma biomarkers, and use the gold particles and near-infrared (NIR) laser light to detect tumors in vivo.4 Pulsing these particles with the NIR beam induced irreversible heat damage to the carcinoma. Subsequent experiments demonstrated this combination of detection and treatment caused a significant increase in lifespan compared with controls.5 These studies were conducted in mouse models, but, by this time, a few first-generation nanoparticletherapeutics had already obtained recognition in the clinical cancer research community as well. Doxil, the liposomal formulation of doxorubicin, had been approved by the FDA in the mid-1990s and had

TABLE 1 Nanoparticles for the Detection and Treatment of Cancer Which Have Been Approved by the FDA or are Presently in Clinical Trials

Product

Type of Nanoparticle/ Drug

Indication

FDA

Company

Doxil

PEGylated liposome/doxorubicin hydrochloride

Ovarian cancer

Approved 11/17/1995 FDA50718

OrthoBiotech

Abraxane

Nanoparticulate albumin/paclitaxel

Various cancers

Approved 1/7/2005 FDA21660

American Pharmaceutical Partners

Cyclosert

Cyclodextrin nanoparticle

Solid tumors

Phase I

Insert Therapeutics

Megace ES

Nanocrystal/megestrol Breast cancer acetate

Approved 7/5/2005 FDA21778

Par Pharmaceutical Companies

INGN-401

Liposomal/FUS1

Lung cancer

Phase I

Introgen

Combidex

Iron oxide

Tumor imaging

Phase III

Advanced Magnetics

Aurimune

Colloidal gold/TNF

Solid tumors

Phase II

CytImmune Sciences

SGT-53

Liposome Tf antibody/p53 gene

Solid tumors

Phase I

SynerGene Therapeutics

Source: http://www.accessdata.fda.gov/scripts/cder/drugsatfda/

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

265

Opinion

www.wiley.com/wires/nanomed

FIGURE 1 | Hypothetical mechanism by which a PEG

coating protects a nanoparticle from recognition by the immune system The left panel shows that the PEG coating prevents the transient interaction and binding of antibodies with proteins on the nanoparticle surface. The right panel shows that the same PEG coating does not interfere with recognition by the cellular receptors.

demonstrated a decreased cardiotoxicity compared with free doxorubicin. Rapamune, a nanocrystal version of sirolimus, was approved in the USA in 2000, and offered an increased solubility because of greater surface area. Abraxane, a nanoscale, albumin-bound form of paclitaxel, was in FDA Phase III trials and was showing decreased toxicity compared with the cremophore-based carrier. The nanoscale component of each of these drugs is a naturally occurring structure or compound—liposomes (Doxil), a crystalline phase (Rapamune), or polymers (Abraxane). Sally’s trial, however, was for a cancer drug attached to a gold nanoparticle—a particle engineered with polymers and biomolecules on its surface to avoid rapid clearance by the reticuloendothelial system (RES) and to target malignant cells, respectively. Such manipulation offers possibilities for pharmacological improvements, since nanoparticle drug-carriers can be tailored to maximize benefits and minimize side effects.

BENEFITS OF NANOTECHNOLOGY Nanoparticles are on the same size scale as receptors, channels, ligands, effectors, and nucleic acids and can be modified or otherwise further engineered to achieve a particular physiological effect, such as increased biocompatibility. Drugs found to be efficacious under in vitro conditions, such as in high-throughput screening studies, are often insoluble—and are rapidly cleared from the bloodstream when injected into animals or people. Hydrophilic molecules such as polyethylene glycol (PEG) can be bound to nanoparticle surfaces, which greatly increase their solubility and biocompatibility.6 Candidate drugs that were previously discarded because of insolubility or high molecular weight can be attached to this nanoparticle 266

‘platform’. Albumin, the most plentiful protein in human serum, turns out to be a natural ‘solvent’ for paclitaxel. Scientists at ABI developed methods for making nanoparticles from the material and loading paclitaxel into them, allowing patients to safely receive 50% more paclitaxel per dose than is possible with the free drug. In addition, experimental data suggest that the albumin nanoparticles (i.e., Abraxane) interact with tumor blood vessel receptors that transport the nanoparticles into tumors. This interactivity may account for the increased levels of paclitaxel seen in tumors treated with Abraxane compared with legacy paclitaxel.7 When attached to an engineered nanoparticle, a drug’s solubility, half-life, and general biocompatibility depend on the tailorable properties of the nanoparticle, rather than the intrinsic properties of the drug itself. Considerable research is now being invested in qualifying nanoparticles as ‘platforms’ for various drugs. Table 1 lists nanotech-based constructs currently in clinical or preclinical development. In addition to improving solubility, nanotech formulations also offer decreased toxicity and increased efficacy because of passive and active targeting.8 Passive targeting exploits the size and surface properties of nanoparticles, which allows them to extravasate through the endothelial wall.9,10 This vascular leakiness and decreased lymphatics enable nanoscale particles to accumulate in tumor tissues, concentrating the attached drug where it is needed, rather than in Sally’s liver. One study reported that up to 8% of the nanoparticle’s total administered dose can accumulate in the tumor by this ‘enhanced permeation and retention’ (EPR) effect.11 In an alternate form of passive targeting, the nanoparticle’s surface chemistry promotes phagocytosis by macrophages and then delivery to sites of inflammation or sentinel lymph nodes. In active

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanoparticle therapeutics

endotoxin contamination All samples shown here were spiked with a known amount (0.4 EU/mL) of UPS grade endotoxin standard. The quality control sample (blue) and the dendrimer sample (yellow) yield signals very close to the theoretical value, and fall within USP limits of ±50%. Quantification of endotoxin in the gold nanoparticle (green) sample is inhibited—the test result is below the acceptable limits, yielding a false-negative endotoxin result. Quantification of the endotoxin in the polymer nanoparticles (burgundy) is enhanced—the test result is above the acceptable limit.

EU/mL

FIGURE 2 | Nanoparticles interfere with the LAL test for

2.0 1.8 1.6 1.4 1.2 1.0 0.8 0.6 0.4 0.2 0

targeting, the nanoparticle includes a molecule that binds to a biomarker at the tissue of interest. In cancer applications, this marker is generally the ligand for an extracellular receptor that is overexpressed on the cell surface compared with nonproliferating cells. Ligands for particular cellular receptors can be attached to a nanoparticle and facilitate active targeting to tissues expressing those receptors.12–14 Examples of these markers include ligands such as folate and transferrin and receptors such as herceptin receptor (HER2). The preferential delivery of nanoparticulate drugs to tumors allows lower dosages to be effective and reduces the adverse side effects of chemotherapeutics. Polymer coatings on the nanotech platform are also thought to contribute to this decreased toxicity. The current model is that polymers, such as PEG, attached to the nanoparticle surface provide a protective shield against recognition and/or opsonization by the immune system (Figure 1). This is an intriguing notion, given that direct PEGylation of recombinant proteins has not lived up to early expectations of mitigating interactions with the immune system (reviewed in Ref.15). Immunotoxicity, in the form of hypersensitivity, is often the biggest hurdle for the use and approval of biotechnology-derived drugs in humans—a $67 billion market in 2006, or roughly 10% of total pharma sales.16 Several nanotech formulations have demonstrated decreased hypersensitivity compared with their legacy counterparts.17 One Abraxane study reported only a mild effect with the nanoparticle formulation, compared with severe hypersensitivity in 2–4% of patients treated with the legacy form of paclitaxel.7 Recombinant-derived protein therapeutics are also prone to aggregation because of their limited solubility, which significantly contributes to their antigenicity (reviewed in Ref.18). Proteins bound to nanoparticle platforms are able to overcome this hydrophobicity barrier, as discussed above. Vo lu me 1, May /Ju n e 2009

+50% Theoretical −50%

Quality control Gold Polymer Dendrimer LAL-free water nanoparticles nanoparticles nanoparticles

CHALLENGES The formulation that was administered to Sally is reddish-black as it is injected into the i.v. tube. The strange coloration is caused by light scattering by the colloidal gold. That is one of the difficulties in characterizing nanoparticles, they often absorb light and interfere with the in vitro methods used to evaluate their physicochemical or immunological properties. Other nanoparticles have catalytic properties and can actually enhance assays that rely on enzymatic reactions, thus generating false-positive results. Assays routine to the preclinical characterization of conventional pharmaceuticals, such as the Limulus amebocyte lysate (LAL) test for endotoxin contamination detection, may yield spurious results when applied to nanoparticle samples (Figure 2). One of the objectives of our lab, the Nanotechnology Characterization Laboratory (NCL), is to develop and qualify a battery of tests that can accurately assess nanoparticle safety and efficacy.19 Early on, we at the NCL attempted to simply adapt off-the-shelf, commercially available kits to evaluate nanotech formulations. That approach was quite na¨ıve—there is very little about nanomedicine that is ‘off-the-shelf’. This complexity can be a deterrent to those who fund drug development, and they are looking for a near-term return on their investment. Yet the colloidal gold formulation given to Sally has beaten the odds and made it to FDA Phase I clinical trials, and without venture-capital funding. Even when well-established methodologies can be used, nanoparticle therapeutics poses a myriad of characterization challenges.20–23 For soluble smallmolecule drugs, measurement of the molecular weight is frequently adequate for identification, but multifunctional nanoparticles have to be characterized much more rigorously, as there are more components that must work in concert with achieving functionality. Meaningful physicochemical characterization of a multifunctional entity such as a nanoparticle includes

 2009 Jo h n Wiley & So n s, In c.

267

Opinion

www.wiley.com/wires/nanomed

Nanoparticle Biocompatibility

Percent Control

(+) 120 100

Low Cytotoxicity (Surface Reactivity)

] ]

80 60

RES Recognition

40 20 0 0.0

0.5

1.0

Solubility

Billary Clearance?

0

Renal Clearance

Zeta potential

xxxx (mg/mL)

(EPR Effect) High

(−) 1 nm

220 nm Size (Rigid Core)

the assessment of the magnitudes of the individual parts, the stoichiometry and connections between the parts, and the chemical stability of those associations (e.g., covalent and van der Waal bonds).20 A thorough characterization of a nanoparticle-based therapeutics includes evaluation of physicochemical properties, sterility and pyrogenicity assessment, biodistribution (ADME or absorption, distribution, metabolism, and excretion), and toxicity characterization—which include both in vitro tests and in vivo animal studies. Each of these tiers of a rational characterization cascade must be tailored so that they are relevant to nanoparticles. One challenge for in vivo studies of multifunctional nanoparticles is tracking the individual constituents of a multicomponent entity. The formulation may be ineffective if the therapeutic (drug) moiety of a particle disassociates from the nanoparticle platform upon administration or during circulation in the blood, is degraded inside the particle, or fails to disassociate in targeted tissue. Premature release of a cytotoxic drug (e.g., a chemotherapeutic) from the platform may result in acute toxicity independent of the phamacokinetics (PK) of the nanoparticle platform. In vivo studies that image/trace only the drug or only the nanoparticle platform may therefore be inadequate for understanding the therapeutic efficacy and toxicity of a multicomponent nanoparticle. Radiolabeled studies that use multiple isotopes (or multiple imaging methods with nonoverlapping signals) to track the various components of a nanoparticle are 268

FIGURE 3 | The physicochemical characteristics of a nanoparticle influence biocompatibility Here we qualitatively show trends in relationships between the independent variables of particle size (neglecting contributions from attached coatings and biologics), particle zeta potential (surface charge), and solubility with the dependent variable of biocompatibility—which includes the route of uptake and clearance (shown in green), cytotoxicity (red), and RES recognition (blue).

more reliable.24 Different phamacokinetic parameters between the components, such as disparate volumes of distribution or half-lives, are indicative of particle instability—and likely to be missed by studies employing single-component radiolabeling. The gold colloid construct administered to Sally is only one of the many nanoparticle types we help develop at the NCL. Our charter is to provide infrastructure support to NCI’s Alliance in Nanotechnology—to assist nanotech researchers as they transition their nanotech constructs into clinical trials. Over the past 3 years, we have characterized liposomes, dendrimers, fullerenes, quantum dots, polymers, and nanoemulsions. Each particle type has its own strengths and limitations, depending on the intended clinical indication. For the most part, all are demonstrating ‘disruptive’ capabilities compared with their small-molecule counterparts. Because of their complexity, however, preclinical characterization has been a rate-limiting phase of the regulatory and commercialization process. To facilitate this assessment, the NCL, in coordination with the FDA and NIST, is developing standardized methods for nanoparticle characterization, methods that work for multiple particle types. In the pharmaceutical industry, small-molecule drugs are generally categorized based on common chemical structures and interaction target sites. For nanoparticle formulations, this categorization, with

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanoparticle therapeutics

respect to biocompatibility, will likely be based on multiple physicochemical properties rather than simply on the class of material. A plethora of data now exists in the scientific literature, as well as characterization data generated by NCL on proprietary formulations, demonstrating that each nanoparticle is unique. Slight changes to a nanoparticle’s size or surface chemistry, for instance, can dramatically influence a physiological response.25–27 At the NCL, we are beginning to ‘bin’ trends in biocompatibility based on size, surface charge, and hydrophobicity, as shown in Figure 3.

FROM BENCH TO BEDSIDE Regulatory review of multifunctional particles, such as the one in Sally’s study, is often the topic of lively discussions at scientific conferences and working groups. One can appreciate that it may not be straightforward to decide whether a multipart multifunctional nanoparticle should be classified, and subsequently regulated, as a device, drug, or biologic. Different centers within the FDA have jurisdiction over these various products, each with its own Guidance documents. The agency may, in fact, consider nanoparticles as combination products—where one center is in ‘lead’ with consultation from its sister centers.28 The lead center is determined by the product’s ‘primary mode of action’. A multifunctional nanoparticle intended for therapeutics, for example, would generally fall under the Center for Drug Evaluation and Research (CDER). In any case, the FDA encourages sponsors to consult with them early in the development phase to facilitate the regulatory review process and to provide course corrections as necessary. In their recently released report from the Nanotechnology Task Force, the agency announced that it intends to generate a nanotechnology-specific guidance document in the near future.29 The empirical data traditionally used to inform Guidance for Industry documents are somewhat sparse for nanomedicine, however, because of the field’s infancy and often disparate methods for nanoparticle evaluation. In addition to preclinical characterization and regulatory hurdles, scale-up often poses a challenge for developers. Proof-of-concept studies, involving milligram quantities of multifunctional nanoparticles,

often make the headlines in scientific journals. But it is quite another thing to produce gram to kilogram quantities for clinical trials, with batch-to-batch consistency. Specific good-manufacturing procedure (GMP) requirements have not yet been established for nanoparticles, and the current requirements cannot readily be applied to nanoparticles. For example, GMP producers familiar with narrow chromatographic peaks for small molecules will have to adjust considerably for multifunctional particles. The particle’s multicomponent nature ensures that 95% purity of each component, assuming a stepwise synthesis, does not equate to 95% purity of the final formulation. The nanoparticle’s ‘polydispersity index’—historically a term referring to a sample’s molecular weight distribution, but now also applied to nanoparticle size distribution—may soon become part of the nanobased drug’s acceptance criteria for quality control of batch release.

CONCLUDING REMARKS As one can likely tell from the tone of this article, my initial skepticism of 4 years ago has been replaced by increasing optimism. Cancer research has already seen improvements in efficacy and decreases in toxicity, despite the relative infancy of the field. Nanomedicine, especially as it applies to cancer therapeutics and diagnostics, has a complex developmental pathway and is not for the faint-hearted investigator—or investor! Yet the payoffs are appreciable. Let us check back on Sally. I mentioned earlier that she was being given a nanoformulation of a drug that had failed in clinical trials 10 years before. The nurses and doctors went in and out of her room and hovered at the nurse’s station; they seemed to be anticipating an acute adverse event, which had been common to the earlier (non-nanotech) trial. Several hours into the treatment they appeared to go back to their rounds and check in on other patients. Did I mention that Sally was given a dose of the drug at a level that would have been lethal, as determined in the failed clinical trial of 10 years ago? Yet, she exhibited no observable side effects for the nanoparticle formulation. Perhaps there really is some science behind this nanotech hype after all.

NOTES This project has been funded in whole or in part with federal funds from the National Cancer Institute, National Institutes of Health, under contract N01-CO-12400. The content of this publication does not necessarily reflect Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

269

Opinion

www.wiley.com/wires/nanomed

the views or policies of the Department of Health and Human Services, nor does the mention of trade names, commercial products, or organizations imply an endorsement by the US Government. The author declares no competing financial interest. The author thanks Dr. Jennifer Hall for her considerable assistance during the preparation of this article.

REFERENCES 1. http://nano.cancer.gov/about alliance/cancer nanotech nology plan.asp (accessed, 2008) 2. Kukowska-Latallo JF, Candido KA, Cao Z, Nigavekar SS, Majoros IJ, et al.: Nanoparticle targeting of anticancer drug improves therapeutic response in animal model of human epithelial cancer. Cancer Res 2005, 65:5317–5324. 3. Bibby DC, Talmadge JE, Dalal MK, Kurz SG, Chytil KM, et al.: Pharmacokinetics and biodistribution of RGD-targeted doxorubicin-loaded nanoparticles in tumor-bearing mice. Int J Pharm 2005, 293:281–290. 4. Hirsch LR, Stafford RJ, Bankson JA, Sershen SR, Rivera B, et al.: Nanoshell-mediated near-infrared thermal therapy of tumors under magnetic resonance guidance. Proc Natl Acad Sci USA 2003, 100:13549–13554. 5. O’Neal DP, Hirsch LR, Halas NJ, Payne JD, West JL. Photo-thermal tumor ablation in mice using near infrared-absorbing nanoparticles. Cancer Lett 2004, 209:171–176. 6. Harris JM, Chess RB. Effect of pegylation on pharmaceuticals. Nat Rev Drug Discov 2003, 2:214–221. 7. Desai N, Trieu V, Yao Z, Louie L, Ci S, et al.: Increased antitumor activity, intratumor paclitaxel concentrations, and endothelial cell transport of cremophor-free, albumin-bound paclitaxel, ABI-007, compared with cremophor-based paclitaxel. Clin Cancer Res 2006, 12:1317–1324.

13. Ravi Kumar M, Hellermann G, Lockey RF, Mohapatra SS. Nanoparticle-mediated gene delivery: state of the art. Expert Opin Biol Ther 2004, 4:1213–1224. 14. Patri AK, Kukowska-Latallo JF, Baker JR Jr. Targeted drug delivery with dendrimers: comparison of the release kinetics of covalently conjugated drug and noncovalent drug inclusion complex. Adv Drug Deliv Rev 2005, 57:2203–2214. 15. Chamberlain P, Mire-Sluis AR. An overview of scientific and regulatory issues for the immunogenicity of biological products. Dev Biol (Basel) 2003, 112:3–11. 16. http://www.researchandmarkets.com/reports/c52103 17. Dobrovolskaia MA, McNeil SE. Immunological properties of engineered nanomaterials. Nat Nano 2007, 2:469–478. 18. Rosenberg AS. Effects of protein aggregates: an immunologic perspective. AAPS J 2006, 8:E501–E507. 19. http://ncl.cancer.gov/ 20. Patri AK, Dobrovolskaia MA, Stern ST, McNeil SE: Preclinical characterization of engineered nanoparticles intended for cancer therapeutics. In: Amiji M: ed. Nanotechnology for Cancer Therapy. Boca Raton, FL: CRC Press/Taylor&Francis; 2006, 105–139.

8. Allen TM, Cullis PR. Drug delivery systems: entering the mainstream. Science 2004, 303:1818–1822.

¨ 21. Oberdorster G, Maynard A, Donaldson K, Castranova V, Fitzpatrick J, et al.: Principles for characterizing the potential human health effects from exposure to nanomaterials: elements of a screening strategy. Part Fibre Toxicol 2005, 2–8.

9. Fang J, Sawa T, Maeda H. Factors and mechanism of ‘‘EPR’’ effect and the enhanced antitumor effects of macromolecular drugs including SMANCS. Adv Exp Med Biol 2003, 519:29–49.

22. Oberdorster G, Oberdorster E, Oberdorster J. Nanotoxicology: an emerging discipline evolving from studies of ultrafine particles. Environ Health Perspect 2005, 113:823–839.

10. Maeda H, Wu J, Sawa T, Matsumura Y, Hori K. Tumor vascular permeability and the EPR effect in macromolecular therapeutics: a review. J Controlled Release 2000, 65:271–284.

23. Powers KW, Brown SC, Krishna VB, Wasdo SC, Moudgil BM, et al.: Research strategies for safety evaluation of nanomaterials. Part VI. Characterization of nanoscale particles for toxicological evaluation. Toxicol Sci 2006, 90:296–303.

11. Kirpotin DB, Drummond DC, Shao Y, Shalaby MR, Hong K, et al.: Antibody targeting of long-circulating lipidic nanoparticles does not increase tumor localization but does increase internalization in animal models. Cancer Res 2006, 66:6732–6740. 12. Sahoo SK, Labhasetwar V. Nanotech approaches to drug delivery and imaging. Drug Discov Today 2003, 8:1112–1120.

270

24. Zolnik BS, Stern ST, Kaiser JM, Heakal Y, Clogston JD, Kester M, McNeil SE: Rapid distribution of liposomal short-chain ceramide in vitro and in vivo. Drug Metab Dispos, 2008. 25. Kobayashi H, Kawamoto S, Jo SK, Bryant HL, Brechbiel MW Jr, et al.: Macromolecular MRI contrast agents with small dendrimers: pharmacokinetic

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanoparticle therapeutics

differences between sizes and cores. Bioconjug Chem 2003, 14:388–394. 26. Malik N, Wiwattanapatapee R, Klopsch R, Lorenz K, Frey H, et al.: Dendrimers: relationship between structure and biocompatibility in vitro, and preliminary studies on the biodistribution of I-125-labelled polyamidoamine dendrimers in vivo. J Controlled Release 2000, 65:133–148.

Vo lu me 1, May /Ju n e 2009

27. Nigavekar SS, Sung LY, Llanes M, El-Jawahri A, Lawrence TS, et al.: H-3 dendrimer nanoparticle organ/tumor distribution. Pharm Res 2004, 21:476–483. 28. http://www.fda.gov/nanotechnology/ 29. http://www.fda.gov/nanotechnology/taskforce/report 2007.pdf

 2009 Jo h n Wiley & So n s, In c.

271

Advanced Review

In vivo visualization of macrophage infiltration and activity in inflammation using magnetic resonance imaging Nicolau Beckmann,1∗ Catherine Cannet,1 Anna Louise Babin,1–3 ´ 2,4 Stefan Zurbruegg,1 Rainer Kneuer1 and Franc¸ois-Xavier Ble, 5 Vincent Dousset Because macrophages play a key role on host defense, visualization of the migration of these cells is of high relevance for both diagnostic purposes and the evaluation of therapeutic interventions. The present article addresses the use of iron oxide and gadolinium-based particles for the noninvasive in vivo detection of macrophage infiltration into inflamed areas by magnetic resonance imaging (MRI). A general introduction on the functions and general characteristics of macrophages is followed by a discussion of some of the agents and acquisition schemes currently used to track the cells in vivo. Attention is then devoted to preclinical and clinical applications in the following disease areas: atherosclerosis and myocardial infarction, stroke, multiple sclerosis, rheumatoid arthritis, and kidney transplantation .  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 272–298

INTRODUCTION

I

nflammation is a complex, highly regulated sequence of events that can be provoked by a variety of stimuli including pathogens, noxious mechanical and chemical agents, and autoimmune responses. The inflammatory response occurs in the vascularized connective tissue, and it includes plasma, circulating cells, blood vessels, and cellular and extracellular components, corresponding with increased microvascular caliber, enhanced vascular permeability, leukocyte recruitment, and release of inflammatory mediators.

∗ Correspondence

to: [email protected]

1 Novartis

Institutes for BioMedical Research Global Imaging Group Forum 1, Novartis Campus WSJ-386.2.09 CH-4056 Basel—Switzerland

2 Respiratory Diseases Department, Novartis Institutes for BioMedical Research, CH-4056 Basel, Switzerland 3 Sackler

Institute of Pulmonary Pharmacology, King’s College, London SE1 1UL, UK 4 Mouse

Imaging Centre, Toronto Centre for Phenogenomics, Toronto, Canada M5T 3H7 5

University Victor Segalen Bordeaux 2, EA 2966 Neurobiology of Myelin Disease Laboratory, CHU de Bordeaux, F-33076 Bordeaux, France DOI: 10.1002/wnan.016

272

Inflammation is the primary process through which the body repairs tissue damage and defends itself against adverse stimuli. In the physiologic condition, the regulated response protects against further injury and clears damaged tissue. In pathologic situations, however, inflammation can result in tissue destruction and lead to organ dysfunction. The process of inflammation is divided into acute and chronic patterns. Acute inflammation is of relatively short duration, lasting for minutes, several hours, or a few days, and its main features are the exudation of fluid and plasma proteins (edema) and the emigration of leukocytes, predominantly neutrophils. Chronic inflammation is of longer duration and is associated histologically with the presence of lymphocytes and macrophages, the proliferation of blood vessels, fibrosis, and tissue necrosis. In inflammation, macrophages have three major functions: antigen presentation, phagocytosis, and immunomodulation through production of various cytokines and growth factors.1–4 The mononuclear phagocyte system (MPS) consists of cells that have a common lineage whose primary function is phagocytosis. Monocytes, the circulating blood-born macrophages, and resident

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

tissue macrophages in their various forms make up the system.2,3 These cells form a system because of their common mesenchymatous origin, similar morphology, and common functions, particularly phagocytosis. Cells of the MPS originate in the bone marrow, circulate in the blood, and mature and become activated in various tissues. The first cell type that enters the peripheral blood after leaving the marrow is incompletely differentiated and is referred to as a monocyte. Upon reaching the extravascular tissues and settling there, monocytes mature and become blood-borne macrophages in the tissues, which are larger phagocytic cells. Macrophages have the potential of being activated, a process that results in increased cell size, more active metabolism, ability to present antigens to other immune cells, increased production of lysosomal enzymes, and greater ability to phagocytose and kill ingested microbes. Resident macrophages, which populate tissues during the embryologic period, are found in all organs and connective tissues and named to designate their location, such as microglial cells in the central nervous system, Kupffer cells in the liver, alveolar macrophages in the lung, and osteoclasts in the bone.1,5 They play a major role in the primary immune defense of tissues. Activation of resident macrophages leads to chemokine secretion that increases the adhesion and crossing of white cells, such as circulating macrophages, lymphocytes and mastocytes, from the blood to the tissue, generating the inflammatory reaction. Due to its key role on host defense, visualization of the migration of cells from the MPS, in particular that of macrophages, is of high relevance for both diagnostic purposes and the evaluation of therapeutic interventions. This requires the development of suitable cell labeling strategies, because many cell types do not freely incorporate cellular contrast agents unless the surface of these agents has been complexed with transfection agents. Such agents include TAT,6,7 poly-L-lysine,8,9 or FuGENE,10 serving as translocation signals and enhancing the uptake of the particles, or by using magnetodendrimers.11–13 For macrophages, however, labeling is straightforward as it is their biological function to internalize foreign particles. Since macrophages readily uptake free particles having no surface modifications, intravenous administration of magnetically labeled nanoparticles that are removed from the circulation by blood-borne cells of the MPS, in particular by monocytes, is a strategy routinely applied for tracking macrophages by magnetic resonance imaging (MRI). Attracted by chemotactic signals, the labeled monocytes migrate Vo lu me 1, May /Ju n e 2009

to sites of tissue inflammation. Activated endothelial cells enable adhesion, rolling, sticking, and finally extravasation of the monocytes into affected tissue, whose contrast appearance in magnetic resonance images may be changed due to the nanoparticles contained in the labeled cells. It cannot be excluded that nanoparticles also extravasate directly through the capillaries, especially in inflamed areas where the vessels are more permeable.14 In this case, the nanoparticles will be taken up by macrophages in the tissue. Uptake efficiency depends on several factors including the size and charge of the nanoparticles, surface area-to-volume ratio, physicochemical and biochemical properties of the coating shell, and the cell type.15–19 Competitive uptake by other cells of the immune system has been found to be negligible, as these cells would have to be harvested and labeled in vitro in order to achieve effective incorporation of marker molecules (see Ref. 20 for a recent review). The present article addresses the use of iron oxide and gadolinium(Gd)-based particles for the noninvasive in vivo detection of macrophage infiltration into inflamed areas by MRI. In the first part, some of the agents and their basic properties are presented, followed by a brief discussion of acquisition schemes. Attention will then be devoted to preclinical and clinical applications from several disease areas.

MAGNETIC RESONANCE IMAGING CONTRAST AGENTS FOR IN VIVO CELL TRACKING Iron Oxide Particles Superparamagnetic MRI contrast agents are nanoparticles consisting of a core of insoluble iron oxides such as magnetite (Fe3 O4 ), maghemite or other ferrites, which can be solubilized by coating with hydrophilic polymers.21,22 Superparamagnetism arises in small magnetic crystallites of less than 10 nm in diameter. The thermal energy is not sufficient to overcome the coupling forces between neighboring atoms resulting in a net overall magnetization for the crystallite, yet it is sufficient to rapidly change the direction of this macroscopic magnetization. In the absence of an external magnetic field, the direction of this magnetization vector fluctuates, and the magnetic field exerted by the particle is zero on the MRI timescale. The application of an external field induces a strong net magnetization on superparamagnetic particles. The core of the iron oxide nanoparticles is of the order of 5–10 nm in diameter.23,24 It is encapsulated by a coating, the composition of which

 2009 Jo h n Wiley & So n s, In c.

273

Advanced Review

www.wiley.com/wires/nanomed

largely governs the biodistribution and pharmacokinetic properties of the particle.18 Nanoparticles are called either mono- or polycrystalline, depending on whether they contain one or several iron oxide crystallite cores. Monocrystalline iron oxide nanoparticles (MION) are, for example, the ultrasmall superparamagnetic iron oxide (USPIO) or cross-linked iron oxide (CLIO) particles, while superparamagnetic iron oxide (SPIO) particles are of polycrystalline nature. Typical coating materials are dextran, dextran derivatives such as carboxydextran, starch, albumin, silicon, or polyacrylamide. Interestingly, for dextrancoated particles there is no experimental evidence revealing covalent dextran-FeO bonds, rather, it is assumed that the interaction between dextran and the SPIO/USPIO surface involves hydrogen bonds between hydroxyl groups of the dextran moiety and the surface oxide hydroxyl groups.25 This noncovalent interaction between the iron oxide core and the coating constitutes a problem when designing target-specific nanoparticles, which may lose their coating and thereby their coupling to the targeting moiety. This issue has been addressed by crosslinking dextran molecules with epichlorhydrin.26 The resulting particles, called CLIOs, can be functionalized, presenting activated amino- or carboxylgroups at their surface for coupling to target-specific moieties such as receptor ligands, antibodies, or oligonucleotides.27 The pharmacokinetic behavior of the nanoparticles is governed by their size and coating. For example, dextran-coated SPIOs (AMI-25) have a plasma halflife of approximately 20 to 30 min in rats, while the corresponding value for USPIOs (AMI-227) is of the order of 5 h at an injected dose of 300 µmol/kg.23 Both SPIOs and USPIOs are removed from the circulation by MPS cells via adsorptive endocytosis28 in a process mediated by the monocyte integrin Mac-1 receptor,29 which is also involved in leukocyte adhesion, complement activation, and phagocytosis. Because of the irregular shape of dextran-coated iron oxide particles composed of polycrystalline aggregates, reported size is dependent on the specific measurement method.24

The broad range of individual particle cluster diameters is expected to affect the rate of uptake. Two factors are mainly responsible for an effective internalization of iron oxide into monocytes: particle size and charge. The phagocytic cellular uptake of iron oxides increases with the particle size. In vitro experiments showed that SPIO (particle size of 50–180) was more efficiently phagocytosed by monocytes than USPIO (size of 20–50 nm) or MION (size of 10–20 nm).28,30 Negatively charged particles are also better incorporated than nonionic ones. For instance, it has been demonstrated that the ionic USPIO, SHU555C, with a carboxydextran coating and a size of 21 nm, is significantly better internalized into monocytes ex vivo as compared with the nonionic ferumoxtran10 with a dextran coating and a comparable size of 20–50 nm.30 Superparamagnetic particles are characterized by high relaxivities R1 and R2 (Table 1). The longitudinal relaxivity R1 decreases with increasing Larmor frequency ω0 according to (1/ω0 )3/2 , high magnetic field strengths are associated with smaller relaxivity values.21 The transverse relaxivity, on the other hand, is independent of ω0 . At typical field strengths used in MRI, iron oxide nanoparticles act as R2 contrast agents, leading to negative signal enhancement when T2 -weighted imaging protocols are used. Nevertheless, the nanoparticles can also be used as R1 contrast agents.22 More recently, SPIOs coated with biocompatible polyacrylamide have been described that exert very high R2 and R2 * relaxivities of 690 to 910 and 620 to 1140 s−1 mM−1 , respectively, while the longitudinal relaxivity R1 is relatively low, between 1.7 and 2.4 s−1 mM−1 .31,32 Micron-sized particles of iron oxide have also been introduced as an alternative with which to label macrophages.33–36 These particles of mean diameter between 0.4 and 1.6 µm can be readily purchased and consist of a styrene-divinyl benzene inert polymer microsphere containing a magnetite core as well as a fluorescent label [e.g., the Envy Green dye (525/565 nm) or the Dragon Green dye (480/520 nm)].36 Ex vivo analyses of murine immune cells using fluorescence spectroscopy showed that

TABLE 1 Magnetic Parameters of Iron Oxide Nanoparticles.24 Name

Type

Coating

Core (nm)

Mean particle size (nm)

R1 (s−1 mM−1 )

USPIO

Dextran

5.4

30

22.7 ± 0.2

Ferumoxtran

AMI-227

Ferumoxide

AMI-25

SPIO

Dextran

4.9

150

Ferumoxsil

AMI-121

SPIO

Silicon

8.4

300

Gd-DTPA

R2 (s−1 mM−1 )

23.7 ± 1.2 3.2 ± 0.9 4.5

53.1 ± 3

107 ± 11

72 ± 12 5.7

Relaxivity values are at 20 MHz. The values for gadolinium diethylenetriaminepentaacetate (Gd-DTPA) are provided for comparison.

274

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

the highest (between 50 and 70%) incorporation of MPIOs was observed in macrophages, followed by B cells. The 1.6 µm particles have been shown to yield a water proton, R2, 130% larger than that of an USPIO preparation.36 Single-cell detection can be achieved in vivo with MRI labeling cells ex vivo with MPIOs.33,37–39 The double (magnetic and fluorescent) labeling of MPIOs is advantageous for in vivo applications using MRI and histological analyses such as fluorescence microscopy. Magnetofluorescent nanoparticles (MFNPs) are internalized by macrophages after systemic administration, and therefore, may serve as agents for the in vivo detection of these proinflammatory phagocytes by multimodal imaging.27,40–42 The dextran-covered nanoparticles of CLIO backbone contain a SPIO core detectable by MRI and are labeled with far-red fluorochromes for fluorescence detection (e.g., VT680; excitation/emission, 673/694 nm). From a clinical prospective, MFNPs are particularly interesting for preoperative MRI followed by intraoperative optical imaging. For example, these particles have been used to delineate gliomas pre- and intraoperatively and to guide surgeons during surgery, as the position of the tumor may change during the operation, and the relationship between normal and abnormal tissues becomes unclear.43 Thus, the use of bimodal contrast agents displays benefits both from the excellent spatial resolution of the MRI and the high sensitivity of the fluorescence imaging.

Gadolinium Complexes Bimodal agents for MRI and fluorescence imaging based on gadolinium Gd complexes also are available for macrophage tracking. Gadofluorine M is an amphiphilic Gd complex with a molecular weight of about 1530 g/mol, a concentration of 200 mmol Gd per liter.44 This agent has been shown to display a high R1 relaxivity (the R1 relaxivity of the labeled cells was 137 mM−1 s−1 at 1.5 T and 80.46 mM−1 s−1 at 3 T) and to be spontaneously phagocytosed by macrophages.45 Ex vivo experiments using human monocytes demonstrated efficient cell labeling after incubation of the cells with 25 mM Gd Gadofluorine M for 12 h, resulting in a maximal uptake of 0.3 fmol Gd/cell without impairment of cell viability. Fluorescence microscopy confirmed internalization of the fluorescent contrast agent by monocytes. Imaging studies showed stable labeling for at least 7 days. Lisy et al.46 demonstrated the feasibility of labeling macrophages ex vivo with bacterial magnetic nanoparticles (magnetosomes; particle diameter: 42 nm) harvested from Magnetospirillum gryphiswaldense and coupled to a Vo lu me 1, May /Ju n e 2009

fluorescent dye (DY-676, [lambda]abs./[lambda]em.= 676 nm/701 nm) for MRI and near-infrared fluorescence (NIRF) imaging. At 1.5 T, the nanoparticles have R1 and R2 relaxivities of 3.2 mM−1 s−1 and 526 mM−1 s−1 , respectively, and fluorochromecoupled magnetosomes exhibit increased fluorescence intensities at wavelengths > 670 nm. Activated synovial macrophages overexpress the folate receptor (FR),47 which is a glycosyl phosphatidylinositol-anchored protein that binds folic acid with a high affinity.48 P866 is an MRI agent consisting of a high relaxivity Gd-chelate (contrastophore) combined with a folic acid moiety (pharmacophore) which binds to the FR and is subsequently internalized into the cytoplasm of the cell via endocytosis.49 The R1 and R2 relaxivities of the agent in water at 60 MHz are 21 s−1 mmol−1 and 30 s−1 mmol−1 , respectively.49 The FR exists in three forms: FR-α and FR-β are upregulated on various tumors of epithelial origin and on activated macrophages in inflammation, respectively, while FR-γ is overexpressed in hematopoietic cells.50 As P866 binds to FR-α and FR-β, the contrast agent provides an improved differentiation between FRpositive neoplastic or inflammatory pathologies and FR-negative normal tissues. Gustafsson et al.51 developed probes to label macrophages through the scavenger receptor-A, a macrophage-specific cell surface protein. Agents were based on maleylated BSA ligands, which have a strong affinity and specificity for the scavenger receptorA,52,53 conjugated to multiple Gd-DOTA groups. In vitro studies showed that the targeted contrast agent had accumulated in macrophages, and solution studies indicated that micromolar concentrations were sufficient to produce MRI contrast.51 Also, specific probes for targeting intraplaque macrophages have been prepared by linking biotinylated monoclonal rat antimouse antibody to murine macrophage scavenger receptor-A (MSR-A) types I and II (anti-CD204) to the surface of Gd-mixed micelles using the biotin-avidinbiotin bridging.54–56

MAGNETIC RESONANCE IMAGING SIGNAL ACQUISITION Usually, T2 /T2 *-weighted sequences are used in combination with the administration of iron oxides, resulting in a negative contrast on MR images. This loss of signal may be problematic in several applications, for instance in the study of atherosclerosis in vessel walls or when locating cells in areas with inherent T2 * effects. As SPIO and USPIO agents also increase the

 2009 Jo h n Wiley & So n s, In c.

275

Advanced Review

www.wiley.com/wires/nanomed

longitudinal proton relaxivity, albeit less sensitive, T1 weighted imaging may also be used for detection.22 In many respects, it is preferable to have positive contrast changes for better delineation of contrast changes. Several approaches have been recently proposed to increase the detection sensitivity of positive contrast generated by SPIO-labeled cells: 1 the use of spectrally selective RF pulses to excite and refocus water off-resonance in regions near the labeled cells;57 2 the use of a steady-state free precession (SSFP) sequence (3D-FIESTA), also known as balanced fast field or trueFISP imaging.58 The detection threshold of superparamagnetic Fe using this technique has been predicted to be in the femtomole range, which is not significantly different from the sensitivity of positron emission tomography (PET) and single photon emission computed tomography (SPECT); 3 the use of inversion-recovery on-resonant water suppression (IRON);59 4 passive detection of paramagnetic markers by using white-marker imaging, in which positive contrast from the paramagnetic oxide is achieved by dephasing the background signal with a slice gradient.60 Thus, the signal is conserved in the region near the agent because this induces a dipole field which compensates for the dephasing gradient; and 5 the application of a Gradient-Echo Acquisition for Superparamagnetic Particles (GRASP) sequence consisting of the modification of a standard gradient-echo sequence with user-controlled z-gradient rephrasing.61

MAGNETIC RESONANCE IMAGING-BASED MACROPHAGE TRACKING IN DISEASE To label macrophages, iron oxide agents are administered systemically about 24 h before the MRI measurements. This long delay ensures that free particles are cleared from the circulation and that all signal alterations arise from particles captured by phagocytotic cells.62,63 It has been shown for applications in the central nervous system (CNS) that delays of 4–6 h failed to demonstrate abnormalities even at high contrast agent doses.63 In principle, USPIOs are the favored agents for in vivo macrophage tracking studies, as their longer circulation times compared to SPIOs theoretically increase the probability of particle uptake by blood monocytes.62 However, in vitro 276

labeling of human monocytes has been shown to be more effectively achieved with SPIO (Ferucarbotran) than with USPIO.28,30,64 Furthermore, it cannot be excluded that a fraction of the iron particles leaks out of vessels, especially in inflamed areas, and is captured by tissue macrophages rather than by blood monocytes. The next examples show that both USPIOs and SPIOs can be successfully used to follow noninvasively by MRI the macrophage infiltration into inflamed regions in several disease areas. The imaging agent, dose, and scanning delay after administration should be carefully chosen for each animal model. It is particularly important to ensure that no free iron particles be found in the inflamed tissue. Therefore, validation of imaging findings by histology is fundamental. Prussian blue staining is often used for the in situ detection of iron uptake, however it lacks in sensitivity. Sequential enhancement of Prussian blue staining by diaminobenzidine and silver/gold impregnation has been proposed as a means to enhance the sensitivity for detecting low amounts of iron.65 Alternatively, electron microscopy or the use of bimodal, magnetic and fluorescent agents followed by histological analyses encompassing, for example, fluorescence microscopy, may help in improving the localization of labeled macrophages.

Atherosclerosis and Myocardial Infarction Acute myocardial infarction occurs when the blood supply to a part of the heart is interrupted, most commonly due to rupture of a vulnerable atherosclerotic plaque. The simplistic view of atherosclerosis as a disorder of pathological lipid deposition has been redefined by the more complex concept of an ongoing inflammatory response. The presence of activated macrophages is an early and consistent marker of the inflammatory nature of atherosclerotic disease.66–68 Macrophages are attracted to lesions where they take up and accumulate oxidized and other modified lowdensity lipoproteins (LDLs) from the bloodstream. Over time, macrophages become filled with lipidladen droplets so as to appear foamy when viewed under a microscope, and they are then termed ‘foam cells’. The formation of foam cells is a key event in atherogenesis, and the massive uptake of modified LDLs by macrophages is believed to be mediated by scavenger receptors.69 It has been demonstrated that USPIO agents are capable of navigating through the interstitial endothelial pores of atherosclerotic plaques.70,71 The USPIOs are then taken up and compartmentalized within functional atherosclerotic plaque macrophages.70,72,73

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

The compartmentalization of the USPIOs within the plaque MPS allows for early detection of nonocclusive atherosclerotic disease by means of USPIO-associated T2 and T2 * shortening effects. Even though USPIOs are also used for bright blood magnetic resonance angiography (MRA),74 in plaques the T2 /T2 * effects are always dominant due to the compartmentalization of the particles. Of note, Hyafil et al.75 have demonstrated recently that macrophages in atherosclerotic plaques of rabbits can be detected with a clinical X-ray computed tomography (CT) scanner after the intravenous injection of a contrast agent formed of iodinated nanoparticles dispersed with surfactant. This contrast agent may become an important adjunct to the clinical evaluation of coronary arteries with CT.

Animal Studies The use of USPIO as a marker of atherosclerosisassociated inflammatory changes in the vessel wall before luminal narrowing is present has been evaluated in hyperlipidemic rabbits.72,76–78 Threedimensional MRA of the thoracic aorta using a conventional paramagnetic contrast agent did not reveal any abnormality. Angiograms acquired during the 5 days following injection of USPIO (ferumoxtran-10) showed increasing signal in the aortic lumen. While the aortic wall of the control rabbits remained smooth and bright, marked susceptibility effects became evident on day 4 within the aortic walls of hyperlipidemic rabbits (Figure 1). Histopathology documented marked iron uptake in macrophages embedded in atherosclerotic plaques of the hyperlipidemic rabbits. Electron microscopy showed multiple cytoplasmic iron particles in macrophages. No such changes were

seen in nonhyperlipidemic control rabbits that had received USPIO, or in hyperlipidemic rabbits that had not received USPIO (Figure 1). The results showed that USPIOs were phagocytosed by macrophages in atherosclerotic plaques of the aortic wall of hyperlipidemic rabbits in a quantity sufficient to cause susceptibility effects detectable by MRI.72,77,78 Annexin V recognizes apoptotic cells by specific molecular interaction with phosphatidylserine, a lipid that is normally sequestered in the inner leaflet of the cell membrane but is translocated to the outer leaflet in apoptotic cells, such as foam cells of atherosclerotic plaques.79 Recent research has shown that phosphatidylserine externalization not only occurs in apoptosis, but is also observed in activated macrophages and stressed cells.80 Annexin V can potentially deliver carried materials (e.g., superparamagnetic contrast agents) to sites containing apoptotic cells, such as high-grade atherosclerotic lesions.81 Smith et al.82 decorated the surface of a SPIO with annexin V. Following the administration of biochemically derivatized annexin V SPIOs to rabbit models of human atherosclerosis, a negative contrast in atheromatous lesions, but not in healthy arteries, was observed by MRI. Targeted SPIOs produced negative contrast at doses that were 2000-fold lower than reported for nonspecific atheroma uptake of untargeted superparamagnetic nanoparticles in plaques in the same animal model. The annexin V SPIO partitioned rapidly and deeply into early apoptotic foamy macrophages in plaques. Contrast in plaque decayed within 2 months.82 In vivo examinations in murine models of atherosclerosis have demonstrated a direct correlation between the MRI contrast enhancement provided

(c)

(a)

(b)

(d)

(e)

(f)

FIGURE 1 | (a, b) Injection of USPIO caused spotted susceptibility effects of aortic wall in a heritable hyperlipidemic rabbit (arrows). (c, d)

Histological analysis with modified Prussian blue staining shows superficial distribution of iron particles in the aortic vessel wall of the same animal (arrows; c, black bar represents 200 µm; magnification 25×; d, black bar represents 50 µm; magnification 100×). (e) Aorta of a non- hyperlipidemic, control rabbit without signal inhomogeneities in MRA, and (f) regular endothelial layer in histological section (black bar represents 200 µm; magnification 25×). Black triangle demarks the intimal and medial layer. FLASH MRA datasets were acquired at day 5 after USPIO injection. (Reprinted, with permission, from Ref. 78. Copyright 2006 Wiley Periodicals, Inc.)

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

277

Advanced Review

www.wiley.com/wires/nanomed

by Gd micelles targeting the macrophage scavenger receptor-A and the macrophage content in plaques.56 Magnetofluorescent nanoparticles (MFNPs) have also been used for multimodal imaging (MRI and NIRF) in murine models of myocardial infarction.40,42,83 Following the complete ligation of the left coronary artery, Sosnovik et al.42 followed the uptake of the MFNP, CLIO-Cy5.5,40 by macrophages in the infarcted myocardium. The agent was injected 48 h after surgery, and MRI and tomographic fluorescence optical imaging were performed 48 h later. MRI of the infarcted mice consistently revealed the presence of left ventricular dilatation with thinning and akinesis of the anterior, lateral, and inferolateral walls of the left ventricle. The presence of negative contrast enhancement, consistent with the accumulation of the magnetic nanoparticles, was clearly seen in the hypocontractile areas of the myocardium. No evidence of nanoparticle accumulation, however, was seen by MRI in the hearts of the sham-operated mice. The accumulation of CLIO-Cy5.5 in the infarcted myocardium could also be seen clearly by fluorescence imaging. The uptake of CLIO-Cy5.5 by macrophages infiltrating the infarcted myocardium was confirmed by fluorescence microscopy and immunohistochemistry.42 In addition, Sosnovik et al.83 used the nanoparticle AnxCLIO-Cy5.584 to study cardiomyocyte apoptosis in a mouse model of transient coronary artery occlusion by MRI and fluorescence optical imaging. In hypokinetic regions of the myocardium, T2 * was significantly lower in the mice given AnxCLIO-Cy5.5, and fluorescence target to background ratio was significantly higher.83

Human Studies Recent investigations established that USPIOenhanced MRI can noninvasively image macrophages in carotid atheromata, with validation in carotid endarterectomy specimens.70,85–89 The technique is being explored as a surrogate endpoint in a randomized clinical trial aiming to evaluate patients with plaques in their arteries to see whether treatment with the cholesterol-lowering drug, atorvastatin, can reduce the amount of inflammation within the artery wall within the first three months of treatment.90 Accumulation of USPIO in macrophages in predominantly ruptured and rupture-prone human atherosclerotic lesions causing signal decreases in the in vivo T2 *-weighted MR images has been demonstrated by Kooi et al.70 Imaging was performed on 11 symptomatic patients scheduled for carotid endarterectomy before, and 24 h and 72 h after administration of USPIO. Of the patients with USPIO uptake, signal changes in the vessel walls 278

divided in quadrants surrounding the lumen were observed in respectively 54 and 35% of all patients’ quadrants for MR images acquired at 24 h and 72 h following USPIO administration. For those quadrants presenting changes, there was a significant signal decrease in regions of interest in the images acquired after 24 h, whereas no significant signal change was found at 72 h after USPIO. Histological and electron microscopical analyses showed the presence of iron particles in macrophages within the plaques in 10 of 11 patients. Histological analysis showed USPIO in 75% of the ruptured and rupture-prone lesions, and in only 7% of the stable lesions. Twenty symptomatic patients presenting risk factors consistent with severe atherosclerotic disease underwent MRI before, and 36 h after USPIO infusion.88 Images were manually segmented into quadrants, and the signal change in each quadrant was calculated after USPIO administration. Patients had a mean symptomatic stenosis of 77% compared with 46% on their asymptomatic side, as measured by conventional angiography. All symptomatic carotid stenoses had inflammation, as evaluated by USPIOenhanced imaging. On the contralateral sides, inflammatory activity was found in 19 (95%) patients. Contralaterally, there were 163 quadrants (57%) with a signal loss after USPIO when compared with 217 quadrants (71%) on the symptomatic side (P = 0.007). This study suggests that investigation of the contralateral side in patients with symptomatic carotid stenosis can demonstrate inflammation in 95% of plaques, despite a mean stenosis of only 46%. Thus, inflammatory activity may be a significant risk factor in asymptomatic disease in patients who have known contralateral symptomatic disease. Up to now the question as to whether asymptomatic patients with 50–69% stenosis would benefit from surgical intervention could not be answered. USPIO imaging may be a useful marker in the selection of patients eligible for surgery, if plaque inflammation can be correlated with the risk of developing clinical symptoms. A prospective study correlating future ischemic events with inflammatory plaque activity will be necessary to confirm this hypothesis. Of note, a USPIO-enhanced MRI case study performed on a patient with symptomatic carotid stenosis showed macrophage location correlating with maximal predicted stresses on the plaque.89 This observation supports the hypothesis that macrophages thin the fibrous cap at points of highest stress, leading to an increased risk of plaque rupture and subsequent stroke. An association between Gd-based contrast enhancement in dynamic contrast-enhanced (DCE) MRI experiments and the degree of macrophage

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

infiltration of plaques has been shown in patients.91,92 A fast enhancement component, vp, has been associated with plaque neovasculature,91 whereas a slower enhancement component, Ktrans, has been associated with vessel permeability.92 These dynamic enhancement parameters are strongly associated with macrophage density in advanced carotid plaques.92 The correlations between Ktrans and histologic markers of inflammation suggest that Ktrans is a quantitative and noninvasive marker of plaque inflammation, which is further supported by the correlation of Ktrans with proinflammatory cardiovascular risk factors, decreased high-density lipoprotein levels, and smoking.

Stroke Local brain inflammation is a pathologic hallmark of ischemic stroke lesions,93–95 being spatiotemporally related to the occurrence of delayed apoptotic cell death.96 Postischemic inflammation is dominated by cells of the innate, nonspecific immune system with resident microglia/brain macrophages and bloodderived monocytes/macrophages being the most important cell types involved. Microglia are activated within minutes of ischemia onset and produces a plethora of inflammatory mediators, which exacerbate tissue damage but may also protect the brain against ischemic and excitotoxic injury.97,98 Blood-derived macrophages are recruited with a delay of at least 24–48 h.99,100 At these later stages, the synthesis of proinflammatory cytokines is already downregulated, whereas various anti-inflammatory and protective factors are progressively expressed.94,101 As inflammatory responses to brain ischemia are heterogeneous with respect to the temporal pattern, the cell types involved, and the pathophysiological implications, noninvasive imaging of macrophage infiltration by MRI may provide potential to gain new insights into these processes.7,102–105 Both T1 and T2 -weighted MRI may be used for detection. However, T2 -weighted imaging has an inherent drawback because the hyperintense edema has the opposite effect on the signal intensity as the USPIO-induced signal attenuation. Hence, the effect of USPIO-enhancement may be underestimated. Therefore T1 -imaging, although less sensitive, might be advantageous because the T1 -relaxation of ischemic tissue remains rather constant.

Animal Studies MRI of macrophage infiltration following the administration of iron oxide nanoparticles has been used in murine106 and rat models of brain infarction.102,107–109 In the first study which used Vo lu me 1, May /Ju n e 2009

USPIO for macrophage imaging in a stroke model, ferumoxtran-10 was injected intravenously into rats at 5 h after permanent middle cerebral artery occlusion (MCAO), and MRI was performed on days 1, 2, 4, and 7.108 On T2 -weighted images, patchy areas of signal loss in the infarctions were found until day 4 and decreased thereafter. Similar findings were obtained in a transient ischemia model in rats.109 The reduction in signal loss could have been due to macrophages leaving the site of inflammation. However, physicochemical processes might also contribute to this effect: after incorporation into an MPS cell the dextran coating of USPIOs is removed enzymatically, and due to this process of phagosomal degradation, the particles are resolved and thus act as paramagnetic and not as superparamagnetic particles upon T2 relaxation. The ability to detect macrophages is thereby decreased. An analogous study was performed in a mouse model of permanent MCAO, in which ferumoxtran10 was administered over 5 h.106 USPIO-enhanced MRI kinetic analysis disclosed an inflammatory response surrounding the ischemic lesion and in the contralateral hemisphere via the corpus callosum. The imaging data collected during the first 36-h postinjury time suggested a spread of USPIO-related signal from ipsi- to contralateral hemisphere. Except for the early phase, imaging data correlated with histochemical analysis showing inflammation remote from the lesion and ingestion of nanoparticles by microglia and/or macrophages. However, the hypointense rim surrounding the ischemic lesion present already 1 h after iron particle injection, and indicating focal iron deposition, could not be related to inflammation as studied histologically. Rather than uptake by macrophages, possible mechanisms for the signal loss observed at the early stage of infarct development could have been the diffusion of iron particles through the damaged blood–brain barrier (BBB) as a direct response of endothelial cells to hypoxia110 with subsequent uptake by activated microglia and/or the intravascular trapping of iron particles within developing thrombi.107 Several studies of photochemically induced ischemia in rats showed accumulation of iron oxide particles in the infarct border zone during subacute stages of lesion development.65,102,103,111 Importantly, iron-related signal changes on MRI were paralleled by macrophage-associated iron deposition detected histochemically on postmortem brain sections. In the study by Kleinschnitz et al.,102 SPIO injection between days 5 and 6 after ischemia, but not at earlier time points, caused typical signal loss on T2 *-weighted images obtained 24 h after injection. Assuming that SPIO was primarily taken up by

 2009 Jo h n Wiley & So n s, In c.

279

Advanced Review

www.wiley.com/wires/nanomed

(a)

(b)

(c)

(d)

(e)

(f)

FIGURE 2 | Coronal T2 *-weighted MR images of the rat brain, 6 days after photothrombosis, with (a–c) and without (d–f) intravenous

application of USPIO 24 h earlier. Representative sections through the most rostral (a, d), central (b, e), and caudal (c, f) levels of the lesion. (Reprinted, with permission, from Ref. 103. Copyright 2006 Wiley Periodicals, Inc.)

circulating phagocytes, this study indicates that infiltration of SPIO-laden macrophages occurred to a significant extent only at the end of the first week after ischemia. This observation is in agreement with earlier results suggesting that the recruitment of hematogenous macrophages occurs in a narrow time interval between days 3 and 6 after ischemia.100,112,113 SPIO-induced signal loss indicated active macrophage transmigration into ischemic infarcts but not their mere presence. Furthermore, SPIO-induced signal loss was independent from the disturbance of the BBB assessed using Gd-DTPA. For the same model, USPIO as contrast agent provided similar results to that obtained using SPIO (Figure 2). Taken together, these findings demonstrate that appropriate timing of contrast agent injection and subsequent MRI is of critical importance for iron oxide particle-based macrophage imaging in brain ischemia. At early stages (within hours of vessel occlusion) signal loss may mainly be caused by intravascular trapping of iron particles within developing thrombi, whereas at later stages (within days) it is attributable to macrophage infiltration.

analysis.116 There was no relationship between USPIO-related signal changes and BBB disruption assessed by Gd-enhanced MRI. Interestingly, two distinct components of USPIO-related signal changes were found, one associated with blood vessels and one representing parenchymal enhancement (Figure 3). Vessel-associated changes appeared as signal loss on T2 /T2 *-weighted images and decreased from the first to second scan after USPIO infusion, most likely reflecting a transient blood pool effect of the contrast agent. Conversely, parenchymal enhancement was mainly evident on T1 -weighted images, increased over time, and matched with the expected distribution of macrophages, suggesting that the USPIOenhancement on T1 -weighted images was related to the infiltration of iron oxide-laden macrophages into the brain.105,114 Similar to these findings, prominent signal changes on T1 -rather than on T2 *-weighted images have been recently reported for patients suffering from multiple sclerosis (MS).117 The reason for this effect is currently unknown.

Human Studies

Microglia and hematogenous macrophages are involved in the brain pathology of MS and the corresponding autoimmune animal model, i.e., experimental allergic encephalomyelitis (EAE). Being the primary sensors of brain pathology, these cells are rapidly recruited to sites of infection, trauma, or autoimmune inflammation. Furthermore, they are

Pilot USPIO-enhanced MRI studies have been performed in stroke patients at the early ischemic phase.105,114,115 USPIO was infused 5 to 6 days after stroke onset, corresponding to the presumed period of hematogenous macrophage recruitment102 and infarct expansion shown in quantitative MRI 280

Multiple Sclerosis

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

(a)

(b)

(c)

(d)

(e)

(f)

FIGURE 3 | USPIO-enhanced MRI of human

brain infarction. Images of a 54-year-old man with infarction of the left middle cerebral artery territory showing the differential development of USPIO-related signal changes over time. Compared with the (a) nonenhanced T1 -weighted image 5 days after stroke, T1 -weighted images (b) 24 and (c) 48 h after USPIO infusion show increasing hyperintense signal enhancement in the periphery of the infarcted parenchyma. (d) Nonenhanced T2 *-weighted images display hyperintense demarcation of the infarcted territory. T2 *-weighted images (e) 24 and (f) 48 h after USPIO infusion show a signal change from hyperintense to hypointense attributable to USPIO perfusion. T2 * signal changes are mainly vessel-associated, and can also be observed in the nonischemic hemisphere, and decrease over time (e, f). (Reprinted, with permission, from Ref. 114. Copyright 2004 Oxford University Press.)

competent presenters of antigens and interact with T cells recruited to the inflamed CNS. They also synthesize a variety of molecules, such as cytokines [tumor necrosis factor interleukins], chemokines, accessory molecules (B7, CD40), complement, cell adhesion glycoproteins (integrins, selectins), reactive oxygen radicals, and neurotrophins, which may exert damaging or protective effects on adjacent axons, myelin, and oligodendrocytes.118,119 Active MS lesions are defined by macrophage activity in the presence of partially demyelinated axons. This has led to the prevailing consensus that a T-cell-dependent, macrophage-mediated, autoimmune attack on constituents in the normal myelin sheath underlies the disease. More recently, the possibility has been raised that oligodendrocyte cell death and associated changes within the myelin sheath initiate local macrophage scavenger activity, with subsequent amplification of the inflammatory response.120

Animal Studies Experimental allergic encephalomyelitis can be induced in mice or rats by active immunization with spinal cord homogenates, whole myelinoligodendrocyte glycoproteins, myelin basic proteins, and proteolipid proteins or fragments thereof.121 Although there are certain differences in disease evolution, all animal models are characterized by an acute phase with strong neurological symptoms including Vo lu me 1, May /Ju n e 2009

paresis of the tail and hind paws. In some models a relapse or a chronic phase can be observed. Macrophages play a pivotal role in the acute inflammatory phase where they can be tracked by MRI following systemic administration of USPIO63,122–124 or SPIO.125 Interestingly, a recent study has demonstrated the feasibility to detect and demarcate inflammatory CNS lesions on sonograms by specific macrophage imaging using SPIO.126 The basis of this approach resides on an increase of the tissue echogenicity because of local iron deposition. Using MRI, macrophage infiltration can usually be observed in the spinal cord, the cerebellum, and the medulla. Repeated administration of USPIO allows visualization of acute infiltrating macrophages and therefore USPIO-enhanced MRI can help to selectively demarcate phases of active inflammation from remission.127 An enhanced cerebrovascular permeability and cellular infiltration mark the onset of early MS lesions. So far, the precise sequence of these events and their role in lesion formation and disease progression remain unknown. The state of the BBB reflects several inflammatory processes including early interaction of leucocytes and endothelial cells, transient opening of tight junctions due to transcytosis, and repair mechanisms.128 Therefore, it needs to be considered that this damage of the endothelial lineage and basement membrane could allow particle efflux from the vascular lumen to brain tissue. Analyses of Gd- and USPIO-enhancing lesions in the rat EAE

 2009 Jo h n Wiley & So n s, In c.

281

Advanced Review

day 9

www.wiley.com/wires/nanomed

day 17

day 14

day 11

FIGURE 4 | In a rat EAE model, BBB

(a) Gd -DTPA

(b) USPIOs 300

3

(c)

200

2 100 1

0

0 9

10

11

12 13 14 15 Days post immunization

model have demonstrated a mismatch between these two parameters.124,127,129 Floris et al.129 provided quantitative evidence that BBB leakage is an early event and precedes massive cellular infiltration in the development of acute EAE (Figure 4). Signal enhancement by the contrast agent, gadolinium diethylenetriaminepentaacetate (Gd-DTPA), reflecting vascular leakage, occurred concomitantly with the onset of neurological signs and was already at a maximal level at this stage of the disease. Immunohistochemical analysis confirmed the presence of serum-derived proteins such as fibrinogen around the brain vessels early in the disease, whereas no cellular infiltrates could be detected. MRI further demonstrated that Gd-DTPA leakage clearly preceded monocyte infiltration as imaged by the contrast agent based on USPIO, which was maximal only during full-blown EAE (Figure 4). Ultrastructural 282

16

17

Mean body weight (g)

Mean clinical score

4

leakage and cellular infiltration as monitored by Gd-DTPA and USPIO contrast imaging. (a) Examples of T1 -weighted gradient-echo difference images, showing the relative increase of the signal intensity after a bolus of 0.5 mM Gd-DTPA (8 min in circulation) at different stages of EAE. Compared with an EAE animal at day 9 postimmunization, signal intensities start to increase in the cerebellum and brain stem from day 11 due to the leakage of the BBB. This leakage is even more pronounced at day 14. Eventually, the BBB integrity is restored at day 17. (b) T2 -weighted images of animals from different brain areas that were acquired 24 h after an intravenous bolus of 600 µmol Fe/kg USPIO. Note that the iron oxide starts to accumulate in the brain (as observed by a reduction of the T2 -weighted signal intensities) from day 11 after immunization. Massive USPIO accumulation was observed at the peak of the disease (day 14), whereas in the recovery phase (day 17), only few USPIOs were present in the CNS. (c) Clinical scores of EAE animals in time. Time course of clinical scores (bar) and body weight (line) in acute EAE of Lewis rats. Lewis rats were immunized with myelin basic protein emulsified in complete Freund’s adjuvant at day Zero. Neurological symptoms were scored daily. Data represent mean ± sem; *P < 0.05. (Reprinted, with permission, from Ref. 129. Copyright 2004 Oxford University Press.)

and immunohistochemical investigations revealed that USPIOs were present in newly infiltrated macrophages within the inflammatory lesions. Data indicate that cerebrovascular leakage and monocytic trafficking into the brain are two distinct processes in the development of inflammatory lesions during MS, which can be monitored online with MRI using USPIOs and Gd-DTPA as contrast agents.124,127,129 As USPIO-enhanced MRI has been shown to predict lesion severity and disease development in relapsing EAE,130 in vivo tracking of macrophages can be used to test anti-inflammatory or neuroprotective drugs. Prophylactic and therapeutic treatment of Lewis rats with the sphingosin-1 phosphate receptor agonist, FTY720, completely abolished infiltration of macrophages.131 In contrast, anti-VLA-4 antibody treatment did not affect macrophage infiltration in the same model, although it reduced lymphocyte

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

infiltration.132 Floris et al.129 have observed an abolishment of macrophage infiltration in EAE rats treated with the immunomodulator Lovastatin.

Human Studies Exploratory studies using ferumoxtran-10 have been performed in MS patients.117,133,134 Twenty-four hours after USPIO, lesions were seen as high signal intensities on T1 -weighted images and low signal intensities on T2 -weighted images. About 30% of the lesions enhanced with Gd-DTPA, a marker of increased permeability of the BBB, were negative for USPIO,117 confirming results obtained in EAE rat models.123 On the other hand, there were patients who showed USPIO-enhanced lesions but no Gdenhanced lesions117 (Figure 5). Spatial and temporal discrepancies between BBB leakage as demonstrated by Gd-enhancement, and cellular infiltration as shown by USPIO-enhancement reported for ferumoxtran10117,134 were also observed in MS patients using the USPIO SHU555C.135 For this agent, most enhancements were detected in areas with an intact BBB, suggesting that cell-specific mechanisms might have been largely responsible for the contrast changes

observed following USPIO administration. Analyses of blood samples revealing incorporation of SHU555C into peripheral blood mononuclear cells substantiated this view.135 In analogy to findings obtained in experimental models or in human stroke,104,105 the visualization of macrophage activity in vivo with USPIO characterizes a distinct cellular and inflammatory event of the dynamic process of MS lesion formation. It has been suggested that USPIO-enhancement may be related with a beneficial aspect of inflammation,135 possibly associated with repair mechanisms of remyelination.138 Also, the results obtained in animal studies that Gd-leakage precedes monocyte infiltration in the development of acute EAE129 might suggest that the discrepancies between Gd- and USPIOenhancements observed in MS patients could be related to differences in the stages of the lesions. Indeed, recent studies performed on a murine EAE model demonstrated that MRI with GdDTPA correlated with immunoglobulin deposits, and activation of astrocytes and microglia.139 However, Gd-DTPA-enhanced lesions were not associated with the more severe pattern of tissue damage, i.e., loss

24 h after USPIO

No agent

Immediately after Gd-DOTA

24 h after USPIO

FIGURE 5 | MRI examinations of two MS patients. For patient 1 (upper row), clear signal changes (enhancement on T1 -weighted, attenuation on T2 *-weighted MRI) were seen on an MS plaque. A mismatch of contrast agents uptake was seen in patient (lower row). While T1 -weighted images showed a large MS lesion that was not enhanced by Gd (arrowheads), the same acquisition performed 24 h after USPIO demonstrated an uptake of the agent at the periphery of the lesion (arrow). According to histological observations,136,137 the number of macrophages at the center of acute MS lesions is usually minor. These observations raise the question whether macrophage tracking using USPIO could be a stronger predictor for MS development than imaging of increased BBB permeability using Gd-DOTA. See Dousset et al.117 for details. (Reprinted, with permission, from Ref. 117 Copyright 2006 American Society of Neuroradiology.)

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

283

Advanced Review

www.wiley.com/wires/nanomed

of axons and myelin, or inflammatory cell infiltrates. These observations indicate the limitations of GdDTPA to evaluate in vivo tissue destruction and to predict the fate of an inflammatory lesion. Moreover, although USPIO-enhancements most probably reflect, primarily, the entrance of labeled monocytes into the CNS, nonspecific leakage of non-macrophageincorporated contrast agent over a temporarily damaged BBB cannot be excluded. Therefore, the fact that the macrophage activity information obtained with USPIO is distinct to the increased BBB permeability seen with Gd deserves further attention, and most likely the complementary character of both approaches is going to provide a true diagnostic improvement in the near future. Careful comparisons between Gd- and USPIO-enhancements in larger cohorts of MS patients are still necessary, bearing in mind the limited capabilities of Gd to assess BBB damage (sensitivity depends on the dose of the agent,140 and subtle BBB changes that do not enhance with Gd may occur in normal-appearing white matter141 ).

Rheumatoid Arthritis Macrophages possess widespread proinflammatory, destructive, and remodeling capabilities that critically contribute to the acute and chronic phases of rheumatoid arthritis (RA).142–144 Activated macrophages constitute key effector cells in RA, a direct correlation existing between the level of macrophage activity and the observed joint inflammation, articular pain, and bone erosion. This relationship may be explained by the fact that activated macrophages secrete multiple potent mediators of inflammation and tissue destruction, including proinflammatory cytokines [e.g., interleukin-1 (IL-1), IL-6, TNF-α], chemokines, prostaglandins, metalloproteinases, and reactive oxygen species.144 Furthermore, activated macrophages are known to participate in antigen presentation, and thereby, they are thought to contribute to the activation and proliferation of antigen-specific T cells and their consequent destructive activities. Therefore, monitoring noninvasively the macrophage infiltration into sites of inflammation may play an important role in RA.

Animal Studies In vivo labeling of macrophages by SPIO particles administered intravenously or intra-articularly has been exploited in murine,145,146 rat,147 and rabbit148 models of rheumatoid arthritis to follow the infiltration of macrophages into inflamed areas in the living animal. 284

In a rat antigen-induced arthritis (AIA) model,149 consisting of systemic immunization with antigen in Freund’s complete adjuvant followed by intraarticular injection of the same soluble antigen (methylated bovine serum albumin), a significant negative correlation was found between the MRI signal intensity in the knee and the histologically determined iron content in macrophages located in the same region of animals that had received SPIO 24 h before image acquisition.147 Starting 2 days postantigen injection, images from arthritic knees exhibited distinctive signal attenuation in the synovium (Figure 6). This signal attenuation was significantly smaller in knees from animals treated with dexamethasone (0.3 mg/kg/day p.o.) and completely absent in contralateral knees that had been challenged with vehicle. These results demonstrate the feasibility of detecting macrophage infiltration into the knee synovium in this AIA model by labeling the cells with SPIO. This readout may have an impact in preclinical studies by shortening the duration of the experimental period and by facilitating the investigation of novel immunomodulatory therapies acting on macrophages. The great advantage of this model is that the contralateral unchallenged knee serves as a reference in the same animal. Also USPIO-enhanced T1 - and T2 -weighted MRI has been used to characterize an AIA model in rats150 on day 5 following intra-articular injection of antigen. At early time points (3 min–2 h) after its intravenous administration, the USPIO agent produced a T1 enhancement of synovial tissue in synovitis, which was of the same magnitude but more prolonged compared to that achieved using Gd-DTPA. In addition, USPIO provided a higher difference between the enhancement patterns of arthritic and normal joints, since normal joints showed no USPIO but some Gd-DTPA-enhancement. At 2 h following USPIO, there were also pronounced signal attenuations in T2 *weighted images which correlated with the presence of macrophages in the inflamed synovium. These data suggest a role for USPIO-enhanced combined T1 - and T2 *-weighted MR scans for addressing vascular and cellular aspects of synovitis in the initial phase of arthritis. Activated synovial macrophages in the inflamed synovium present key effector cells in RA.143,144 The quantity of activated macrophages in human arthritic joints has been positively related with the grade of articular destruction.144 The activated synovial macrophages overexpress the folate receptor, while normal noninflamed joint components do not.47 In a rat AIA model, Saborowski et al.49 tested the folate receptor-targeted Gd-chelate, P866,

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

(a)

(b)

Day 2

Day 4

2 mm

Day 9

Day 16

(d) 110 100 90 80 70 60 50 40 30 20 10 0

# macrophages (mm−2)

Relative MRI signal (%)

(c)

#

Challenge ##

Vehicle Dexamethasone Control base 0

2

4

9

10

Time after challenge (days)

300 250 200 150 100

P < 0.001

50 0

Vehicle

Dexamethasone

FIGURE 6 | (a) Coronal sections extracted from 3D images of the right knee of a placebo-treated rat, acquired at different time points after mBSA

injection into the same knee. SPIO had been administered 24 h before each measurement. The arrows highlight the signal attenuation in areas corresponding to synovial fluid effusion. (b) Coronal slice through the knee of a vehicle-treated rat, 16 days after antigen administration. Berlin-blue-Giemsa-stained histological images from the same knee. Arrows indicate a massive presence of iron-loaded macrophages. (c) Course of MRI signal intensities (mean ± sem, N = 5 for each group of knees) relative to prechallenge values. Control knees refer to contralateral, vehicle-treated knees, from the same rats that had received antigen on the ipsilateral joints. The levels of significance, *** P < 0.001, and # 0.01 < P < 0.05, ## 0.001 < P < 0.01, correspond to ANOVA comparisons carried out for antigen-challenged knees from placebo- and dexamethasone-treated animals, respectively. (d) Macrophage numbers (mean ± sem, N = 5) determined histologically on day 16 after challenge, in knees from untreated and dexamethasone-treated rats. The level of significance refers to a t-test comparison between both groups of animals. (Modified, with permission, from Ref. 147. Copyright 2003 Wiley Periodicals, Inc.)

and compared the MRI results with the nonfolatereceptor-targeted P866 analog, P1001, or with GdDOTA. All three contrast agents showed an initial perfusion effect with significantly higher contrast agent uptake in arthritic compared to normal knees. In addition, P866, but not P1001 or Gd-DOTA, showed a prolonged enhancement of the synovitis. Histopathology confirmed the presence of folate receptors in the inflamed joints, but not in normal joints,49 suggesting a specific accumulation of the folate receptor-targeted Gd-chelate, P866, in this arthritis model. In addition to MRI, optical imaging techniques and scintigraphy have shown potential in assessing Vo lu me 1, May /Ju n e 2009

macrophage infiltration in RA models. Early signs of experimental arthritis have also been measured in a murine model of AIA, in which the target of the near-infrared fluorescence (NIRF) probe was the F4/80 antigen present on the surface of macrophages infiltrating the inflamed synovial membrane.151 The feasibility of targeting the folate receptor in macrophages has been demonstrated in a murine model of lipopolysaccharide-induced arthritis by NIRF imaging in combination with the administration of a folate-targeting optical probe,152 and in a dog model by scintigraphy following injection of a 99m Tc-based imaging agent.153

 2009 Jo h n Wiley & So n s, In c.

285

Advanced Review

www.wiley.com/wires/nanomed

Human Studies

contrast and tumor growth, and (3) tumor-tobackground contrast and Perl’s staining score. The registered MR enhancement was not influenced by the apoptotic index and by the vascular density in these experimental xenografts.159 Valable et al.160 showed that murine monocytes/macrophages can be labeled simply and efficiently with large, green-fluorescent, MPIOs. Neither size nor proliferation rate of the cells is significantly affected by this labeling. The labeled monocytes/macrophages, administered intravenously to rats which had developed a glioma following stereotactic injection of C6 cells, targeted the brain tumors, a process that could be monitored noninvasively using T2 *-weighted MRI.160 Overall, the results of this study suggest that the use of monocytes/macrophages may be envisaged in the clinic for vectorizing therapeutic agents towards gliomas. The pathologic basis for enhancement patterns of hyperplastic and tumor lymph nodes shown by MR lymphography after intravenous injection of USPIO nanoparticles (AMI-227) has been studied in rats that had received Freund’s adjuvant subcutaneously into the base of the tail and the pad of the right paw.161 In this model, primary arthritis appeared by the seventh day, followed by a spontaneous recovery. Secondary arthritis developed within the fourth week in the hip and knee, and hyperplastic nodes could be easily perceived after the sixth week in the inguinal regions, sometimes on both sides. Measurements performed at 1.5 T showed that following intravenous injection of AMI-227, a decrease in signal intensity indicated active uptake of particles into macrophages, whereas an increase in signal intensity indicated altered capillary permeability in the tumor.161 These findings suggested the ability of the technique to distinguish benign and malignant enlarged lymph nodes. Differences in AMI227 uptake at MRI helping to differentiate tumorbearing from nontumor-bearing nodes were also observed in rats inoculated with cell suspensions of R3327-MATLyLu rat prostate carcinoma or complete Freund adjuvant to generate ipsilateral popliteal lymph node metastases or lymphadenitis.162

To our knowledge, no MRI study has been published until now on the use of nanoparticles for characterizing macrophage infiltration in arthritis patients. As discussed above, several preclinical studies involving the administration of iron oxide or Gd-containing nanoparticles have shown that MRI provides a valuable image marker of macrophage content within the inflamed synovium. Prior to the application of this approach in clinical trials, it will be necessary to validate the nanoparticle uptake in macrophages of RA patients, to determine the responsiveness of the technique to changes in disease activity, and to demonstrate reproducibility within a heterogeneous population. Such efforts would be worthy since with the development of diseasemodifying antirheumatic drugs directed towards suppressing synovial macrophage activity at an early phase of the disease,154 the ability to noninvasively detect and monitor synovial macrophage content has high clinical and research importance.

Cancer Cancer cells secrete a variety of chemoattractants that attract macrophages and cause them to accumulate in the tumor tissue wherein the macrophage becomes a tumor-associated macrophage (TAM) (see Refs 155,156 for recent reviews). Tumors recruit macrophages which may be classified as inflammatory type 1 macrophages (M1) which have antitumor activity or M2 macrophages which are proangiogenic and stimulate tumor growth. Whether TAMs show protumorigenesis or antitumor activity depends on their interaction with cancer cells, other stromal cells, and the tumor microenvironment. The prognosis associated with TAMs is dependent on tumor type, but in breast cancer157 and prostate cancer,158 TAM accumulation has been linked to decreased survival. This fact obviously brings an important perspective for macrophage tracking as a potential diagnostic tool in cancer.

Animal Studies In a murine model in which fragments of human malignant glioma were orthotopically xenografted into the brain of nude mice, Kremer et al.159 showed that tumor contrast enhancement in T1 weighted spin-echo MR images 24 h after intravenous ferumoxtran-10 administration was well correlated to a tumor proliferative index and tumor growth, suggesting that it could be used as an indirect marker of tumor proliferation. A good relationship was observed between (1) tumor-to-background contrast and proliferative index, (2) tumor-to-background 286

Human Studies In the clinics, lymph node metastases have been traditionally diagnosed with CT and ultrasonography. However, the accuracy of such diagnoses, based on size and shape criteria, has not been adequate. The clinical value of MR lymphography using ferumoxtran-10 has been recently reviewed by Barentsz et al.163 In principle, the technique allows the detection of small and otherwise undetectable

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

lymph node metastases in patients with prostate cancer. This has an important clinical impact, as the diagnosis is more precise and less invasive, thus reducing morbidity and healthcare costs. In patients with a suspicion for a recurrence, this technique may show metastatic nodes when they are still small, thus allowing earlier adequate therapy. Identifying small pathologic nodes facilitates radiation therapy. This may result in an increased dose on the malignant nodes and a decreased dose, with reduced side effects, on normal tissues. However, thorough knowledge of sequence parameters and planes, lymph node anatomy, and appearance of normal and abnormal nodes, is essential when using MR lymphography. As important applications, the value of MR lymphography based on USPIO administration has been demonstrated in patients with retroperitoneal and pelvic cancer,164 in gastric cancer patients,165 and in esophageal cancer patients.166 In malignant brain tumors, ferumoxtran-10 has been demonstrated to reveal areas of enhancement, even with a 0.15 T intraoperative MR, that did not enhance with Gd.167 Ferumoxtran-enhancing lesions had persistent increased T1 signal intensity for 2–5 days, which may provide advantages over Gd for postoperative imaging. Histochemistry for iron showed uptake of ferumoxtran in reactive cells (astrocytes and macrophages) rather than tumor cells.167 More recent results indicate that USPIO agents will not replace Gd in the workup of patients with brain tumors.168 USPIOs offer complementary information and may help to differentiate between brain tumors and areas of radiation necrosis.

Kidney Transplantation Transplantation is nowadays the preferred and accepted treatment option in end-stage organ disease. Since the introduction of calcineurin inhibitors as immunosuppressive therapy, acute rejection can be well managed. However, chronic rejection remains the main complication, and its treatment represents a major challenge. Many methods have been proposed to minimize the immunosuppressive requirements following allotransplantation. Major reasons for these efforts are the clinical risks associated with immunosuppression as well as the specific safety concerns of the molecules used. Before these regimens can be developed for clinical application, they require validation in animal models that are reasonably predictive of their performance in humans.169 The gold standard for rejection surveillance after organ transplantation is biopsy, which is invasive and prone to sampling errors. Noninvasive alternatives Vo lu me 1, May /Ju n e 2009

that can complement biopsy are highly desirable. Macrophage tracking by MRI has been proposed as a noninvasive means to monitor the process of graft rejection. Despite the fact that macrophage tracking by MRI has been applied to heart and lung transplantation as well,170 including the use of MPIOs,35 in this contribution we are going to focus our attention on kidney transplants. Infiltration of predominantly recipient-derived lymphocytes and macrophages is a prominent feature of allograft rejection.171 Studies in animal models and in humans have emphasized the presence of macrophages during acute renal allograft rejection,171,172 and they support the hypothesis that macrophage-derived inflammation is a cofactor for chronic allograft nephropathy,173–175 with monocytes/macrophages and T cells being the predominant graft-invading cells of renal allografts with chronic rejection.176–178

Animal Studies Orthotopic kidney transplantation models are adopted in rodents to study renal allograft rejection.179 One of the recipient’s kidneys is nephrectomized and replaced with a kidney from a syngeneic animal or from an allogeneic animal. The graft may have a life-supporting function, in which case the contralateral kidney from the recipient animal is also removed. However, the recipient’s contralateral kidney may as well remain in place, in which case the entire rejection process can be followed with minimal physiological alterations. Detection of macrophage infiltration into allografts by MRI in combination with the administration of iron oxide nanoparticles has been applied to characterize noninvasively the graft rejection status in several rat models of transplantation. Macrophage infiltration during the acute rejection process of kidney grafts has been demonstrated in the Dark Agouti (DA) to the Brown Norway (BN) model.180 Administration of USPIO particles at day 4 post-transplantation led one day later to distinct signal attenuation in the cortex of allogeneic kidneys. Immunohistological staining for ED1+ macrophages and CD4+ and CD8+ T cells in allogeneic transplanted kidneys indicated the accumulation of these immune cells as acute rejection occurred. Morphological studies by electron microscopy confirmed the existence of iron particles inside the lysosomes of macrophages of rejecting kidneys, while Prussian blue staining detected the presence of iron plaques in macrophages. However, no signal reduction was observed in isografts

 2009 Jo h n Wiley & So n s, In c.

287

Advanced Review

www.wiley.com/wires/nanomed

and allografts of recipients receiving triple immunosuppressant treatment with daily subcutaneous injections of Methylprednisolone (2 mg/kg), Rapamycin (1 mg/kg), and Cyclosporine A (CsA) (5 mg/kg).180 In the chronic Fisher 344 to Lewis model, starting 12 weeks post-transplantation, MR images from grafts of untreated recipients that had received SPIO contrast agent (mean size of particles, 150 nm) exhibited distinctive signal attenuation in the cortex181 (Figure 7). Animals treated with CsA (Neoral 1.5 mg/kg/day p.o. by gavage for 10 days after transplantation) to prevent acute rejection showed a signal attenuation in the cortex at 32 weeks posttransplantation, while kidneys from rats treated additionally with everolimus (Certican 1.25 mg/kg/day p.o. by gavage until the end of the study), a rapamycinderivative, had no changes in anatomical appearance. A significant negative correlation was found between the MRI cortical signal intensity and the histologically determined iron content in macrophages located in the cortex. Moreover, a very strong and highly significant negative correlation was found between the MRI signal in the cortex and the rejection scores according to the Banff classification,182 suggesting that this method might be considered as alternative to histological evaluation of kidney biopsies. Renography revealed a significantly reduced functionality of the kidneys of untreated controls 32 weeks after transplantation, while no significant changes in perfusion were observed in any group of rats. Thus by labeling macrophages with SPIO, graft nephropathy was detected by MRI significantly earlier than changes in graft function occurred. Moreover, creatinine levels in the blood and in the urine, the traditional biochemical marker of rejection, remained unchanged182 (Figure 7). These results indicate that monitoring of macrophage infiltration by MRI provides an early means for detecting kidney graft rejection.

Human Studies In a pilot study, Hauger et al.183 evaluated the detection and characterization of macrophage infiltration in native and transplanted kidneys using USPIO. Twelve patients scheduled for renal biopsy for acute or rapidly progressive renal failure (n = 7) or renal graft rejection (n = 5) completed the study. MRI was performed before, immediately after, and 72 h following intravenous injection of USPIO. Histological examination showed cortical macrophage infiltration in four patients (>5 macrophages/mm2 ), two in native kidneys (proliferative extracapillary glomerulonephritis) and two in transplants (acute rejection). These patients showed a 33% mean cortical signal loss on T2 *-weighted images. In the remaining eight 288

patients (showing less than 5 macrophages/mm2 ), there was no cortical signal loss. However, in three of these, presenting with ischemic acute tubular necrosis, a strong signal drop was found in the medulla exclusively. These preliminary results indicate that USPIO-enhanced MRI can demonstrate infiltration of the kidneys by macrophages both in native and transplanted kidneys and may help to differentiate between kidney diseases.

CONCLUSIONS The examples from several disease areas provided above illustrate the usefulness of in vivo tracking of macrophages using MRI in combination with the administration of iron oxide particles, both from a diagnostic and a therapeutic perspective. Currently most applications are preclinical in small animals. This does not limit the importance of the approach: as time courses are easily assessable, it is well suited to be included in experimental studies in animal models of disease addressing inflammatory aspects. As an immediate consequence, after proper validation, the technique can be easily incorporated into drug testing in such in vivo models, a fundamental activity in the framework of pharmaceutical research and development.184–186 Of important note, as mentioned earlier at several points, leakage of the agents out of the vascular bed cannot be excluded, in particular as vessels tend to be more permeable in inflamed areas, and BBB breakdown and vascular endothelial damage may occur in case of CNS injury. For each model, validation studies should ideally include control animals (noninjured and/or macrophage depleted animals), an analysis of the BBB (for CNS models), and quantitative histology at several time points after SPIO or USPIO administration. It is also very important to study the signal changes as a function of the dose of the agent and/or following anti-inflammatory treatment, and verify histologically whether all detected iron is contained in macrophages. Correlations between MRI signal intensities and histological quantification of iron content are essential. From a pragmatic point of view, even if iron oxide nanoparticles administered intravenously leak out of the vessels and are taken up by macrophages in the inflamed tissue rather than by monocytes in the blood, it is very plausible that the contrast changes detected by MRI reflect primarily inflammatory processes as long as no free iron nanoparticles are present. This view holds if the number of resident macrophages (present in the noninflamed tissue) is much lower than that of blood-borne macrophages (attracted to inflamed

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

Vehicle n = 5 Neoral (10 d) Neoral (10 d) Certican

1.4

(c) 1.6 1.4

Iron content (%)

(b)

Cortical signal intensity (a.u)

(a)

1.2 1.0 0.8

Vehicle

1.2 1.0

Neoral (10 d)

0.8 0.6 Neoral (10 d) Certican

0.4 0.2 0.0

0.6 4

8

12

24

32

0.6

Time after transplantation (weeks)

(d)

1.0

1.2

1.4

Blood Unine

100

Creatinine (µmol/l)

0.8

Cortical MRI signal intensity (a.u)

80 60

20 16 12 8 4 4

8

12

16

24

32

Time after transplantation (weeks)

FIGURE 7 | (a) Gradient-echo images of a Fisher kidney transplanted into a Lewis rat, acquired at several time points with respect to

transplantation. The recipient received SPIO 4 weeks prior to each measurement. Images were acquired without respiratory gating. On the right side, histology of Perl’s Prussian blue reaction at week 32 demonstrating iron-loaded macrophages in the kidney cortex. (b) Cortical MRI signal intensity (means ± sem) in grafts from vehicle-treated recipients, from recipients treated with Sandimmun Neoral (1.5 mg/kg/day p.o. for 10 days) and from recipients treated with Neoral (1.5 mg/kg/day p.o. for 10 days) followed by Certican (1.25 mg/kg/day) for the remainder of the study. (c) A significant negative correlation (r = −0.86, P < 0.0001) was found at week 32 between the iron content determined histologically and the MRI signal intensity. (d) The creatinine in the blood and the urine for the groups of recipients remained unchanged during the experimental time. See Refs 181,182 for more details. (Modified, with permission, from Ref. 181. Copyright 2003 Wiley Periodicals, Inc.)

tissue), a situation that occurs in most organs except the lung. At least for experimental studies in animals, the transfusion of ex vivo iron oxide labeled monocytes could be a plausible means to avoid the influence of multiple pathophysiological events in the enhancement patterns detected by MRI following intravascular injection of iron oxide nanoparticles. This approach has been recently applied to a rat model of experimentally induced photothrombosis.111 Results showed not only the feasibility of longitudinal tracking of ex vivo SPIO-labeled monocytes, but suggested that contrast enhancement after injection of free USPIO did not primarily represent signals from peripherally labeled monocytes that migrated toward the inflammatory lesion. Vo lu me 1, May /Ju n e 2009

Very encouraging results have been recently provided by clinical pilot studies as well. Nonetheless, many steps still need to be completed before this approach can be routinely applied in the clinics, e.g., to validate the particle uptake in macrophages of patients, to determine the responsiveness of the technique to changes in disease activity, and to assess reproducibility in a heterogeneous population. Measurements performed in animals may help to guide corresponding studies in humans. While dedicated animal systems are mostly adopted, the capabilities of clinical scanners to generate highresolution images are being improved. On the other hand, small animal MR scanners based on clinical consoles are also being introduced. Such systems

 2009 Jo h n Wiley & So n s, In c.

289

Advanced Review

www.wiley.com/wires/nanomed

ensure easy protocol transfer from and to clinical scanners, and should therefore facilitate translational studies concerning macrophage tracking. Despite such important technical developments, a major issue concerning translational research in this area is the fact that most of the MRI agents used to label macrophages are only experimental. Ferumoxides are approved by the US Food and Drug Administration (FDA), but USPIO nanoparticles, such as ferumoxtran-10 and ferumoxytol, are still investigational. Ferumoxtran-10 has completed phase 3 clinical trials187,188 and has been shown to be safe and to improve diagnostic performance. On the other hand, most macromolecular Gd chelates are in an early, preclinical research stage and the time required to reach clinical applicability is considerable. Furthermore, toxicity concerns such as the prolonged retention of heavily Gd-loaded molecules and of nanoparticles within the organism must be disproved.189 Recent in vitro evidence suggests that ferumoxtran-10 is not toxic to human monocytes and macrophages, does not activate them to produce proinflammatory cytokines or superoxide anions, is not chemotactic, and does not interfere with Fc-receptor-mediated phagocytosis.190 Furthermore, extremely high intracellular ferumoxtran-10 concentrations had only slight or no effects on these key activities. On the other hand, Siglienti et al.191 tested the cytokine production of rat and mouse macrophages in vitro and found that internalization of SPIO/USPIO shifted macrophages towards an antiinflammatory, less responsive phenotype by enhancing IL-10 and inhibiting TNF-α production. During macrophage interaction with T cells, IL-12p40 secretion was inhibited. Based on these findings, potential immunomodulatory effects of SPIO/USPIO particles in vivo warrant further investigation. As already mentioned in the Rheumatoid Arthritis section, nuclear imaging techniques are also being used to image macrophage infiltration and activity in inflamed areas.192–197 Obvious advantages of these techniques are their high sensitivity, enabling the

use of nanoparticles at lower concentrations than MRI permits, and the possibility to develop probes to target specifically molecular markers related to macrophage activity and, therefore, to inflammation. For instance, based on the fact that macrophages, and also reactive microglia, express the peripheraltype benzodiazepine receptor (PBR),198,199 radioligands have been developed to image this receptor in the inflamed brain.193,196,197,200 One of the radioligands is [11C](R)-PK11195, which has been shown to bind specifically to the peripheral PBR enriched in activated brain macrophages.201 A 99m Tc-histidinefolate targeting the FR, which is upregulated on various tumors of epithelial origin and on activated macrophages in inflammation,50 has been successfully used recently in SPECT/CT investigations involving (FR-positive) KB tumor-bearing mice.194 Nahrendorf et al.195 developed a dextranated and DTPA-modified magnetofluorescent 20-nm-nanoparticle labeled with the PET tracer 64 Cu to yield a PET, MR, and an optically detectable imaging agent. In atherosclerotic mice, the trimodal nanoparticle was shown to directly detect macrophages in atherosclerotic plaques. Advantages of the approach include the improved sensitivity, the direct correlation of PET signals with CD68 (a macrophage marker that correlates well with lesion severity and therapeutic modulations202 ), the ability to readily quantify the PET signal, perform wholebody vascular surveys, spatially localize and follow the reporter by microscopy, and the clinical translatability of the agent given similarities to MRI probes in clinical trials. Overall, considering the central role played by macrophages in disease, from CNS disorders to diabetes, the efforts briefly addressed in this contribution are definitely worth pursuing. Noninvasive visualization of the migration of macrophages is relevant for research and diagnostic purposes, as well as for the evaluation of therapeutic interventions. The biomedical research community is already profiting from this important tool, and medical practice and public health will also benefit from it in the near future.

REFERENCES 1. Barron KD. The microglial cell. A historical review. J Neurol Sci 1995, 134(suppl.):57–68.

3. Hume DA. The mononuclear phagocyte system. Curr Opin Immunol 2006, 18(1):49–53.

2. Hume DA, Ross IL, Himes SR, Sasmono RT, Wells CA, et al. The mononuclear phagocyte system revisited. J Leukoc Biol 2002, 72(4):621–627.

4. Fujiwara N, Kobayashi K. Macrophages in inflammation. Curr Drug Targets Inflamm Allergy 2005, 4(3):281–286.

290

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

5. Gordon S. The macrophage. Bioessays 1995, 17(11):977–986. 6. Dodd CH, Hsu HC, Chu WJ, Yang P, Zhang HG, et al. Normal T-cell response and in vivo magnetic resonance imaging of T cells loaded with HIV transactivator-peptide-derived superparamagnetic nanoparticles. J Immunol Methods 2001, 256(1–2):89–105. 7. Wunderbaldinger P, Josephson L, Weissleder R. Tat peptide directs enhanced clearance and hepatic permeability of magnetic nanoparticles. Bioconjug Chem 2002, 13(2):264–268. 8. Frank JA, Zywicke H, Jordan EK, Mitchell J, Lewis BK, et al. Magnetic intracellular labeling of mammalian cells by combining (FDA-approved) superparamagnetic iron oxide MR contrast agents and commonly used transfection agents. Acad Radiol 2002, 9(suppl 2):S484–S487. 9. Frank JA, Miller BR, Arbab AS, Zywicke HA, Jordan EK, et al. Clinically applicable labeling of mammalian and stem cells by combining superparamagnetic iron oxides and transfection agents. Radiology 2003, 228(2):480–487. 10. Arbab AS, Yocum GT, Wilson LB, Parwana A, Jordan EK, et al. Comparison of transfection agents in forming complexes with ferumoxides, cell labeling efficiency, and cellular viability. Mol Imaging 2004, 3(1):24–32. 11. Bulte JW, Douglas T, Witwer B, Zhang SC, Strable E, et al. Magnetodendrimers allow endosomal magnetic labeling and in vivo tracking of stem cells. Nat Biotechnol 2001, 19(12):1141–1147. 12. Bulte JW, Douglas T, Witwer B, Zhang SC, Lewis BK, et al. Monitoring stem cell therapy in vivo using magnetodendrimers as a new class of cellular MR contrast agents. Acad Radiol 2002, 9(suppl 2):S332–S335. 13. Bulte JW, Kraitchman DL. Iron oxide MR contrast agents for molecular and cellular imaging. NMR Biomed 2004, 17(7):484–499. 14. Renkin EM. Multiple pathways of capillary permeability. Circ Res 1977, 41(6):735–743. 15. Pratten MK, Lloyd JB. Pinocytosis and phagocytosis: the effect of size of a particle substrate on its mode of capture by rat peritoneal macrophages cultured in vitro. Biochim Biophys Acta 1986, 881(3):307–313.

18. Cengelli F, Maysinger D, Tschudi-Monnet F, Montet X, Corot C, et al. Interaction of functionalized superparamagnetic iron oxide nanoparticles with brain structures. J Pharmacol Exp Ther 2006, 318(1):108–116. 19. Mowat P, Franconi F, Chapon C, Lemaire L, Dorat J, et al. Evaluating SPIO-labelled cell MR efficiency by three-dimensional quantitative T2* MRI. NMR Biomed 2007, 20(1):21–27. 20. Dousset V, Tourdias T, Brochet B, Boiziau C, Petry K. How to trace stem cells for MRI evaluation? J Neurol Sci 2008, 265(1–2):122–126. 21. Renshaw PF, Owen CS, McLaughlin AC, Frey TG, Leigh JS. Ferromagnetic contrast agents - a new approach. Magn Reson Med 1986, 3(2):217–225. 22. Chambon C, Clement O, Le Blanche A, SchoumanClaeys E, Frija G. Superparamagnetic iron oxides as positive MR contrast agents: in vitro and in vivo evidence. Magn Reson Imaging 1993, 11(4):509–519. 23. Jung CW. Surface properties of superparamagnetic iron oxide MR contrast agents: ferumoxides, ferumoxtran, ferumoxsil. Magn Reson Imaging 1995, 13(5):675–691. 24. Jung CW, Jacobs P. Physical and chemical properties of superparamagnetic iron oxide MR contrast agents: ferumoxides, ferumoxtran, ferumoxsil. Magn Reson Imaging 1995, 13(5):661–674. 25. Bellin M-F, Roy C, Kinkel K, Thoumas D, Zaim S, et al. Lymph node metastases: safety and effectiveness of MR imaging with ultrasmall superparamagnetic iron oxide particles - initial clinical experience. Radiology 1998, 207(3):799–808. 26. Josephson L, Tung CH, Moore A, Weissleder R. Highefficiency intracellular magnetic labeling with novel superparamagnetic-tat peptide conjugates. Bioconjug Chem 1999, 10(2):186–191. 27. McCarthy JR, Kelly KA, Sun EY, Weissleder R. Targeted delivery of multifunctional magnetic nanoparticles. Nanomedicine 2007, 2(2):153–167. 28. Raynal I, Prigent P, Peyramaure S, Najid A, Rebuzzi C, et al. Macrophage endocytosis of superparamagnetic iron oxide nanoparticles - mechanisms and comparison of ferumoxides and ferumoxtran-10. Invest Radiol 2004, 39(1):56–63.

16. Papisov MI, Bogdanov A Jr, Schaffer B, Nossiff N, Shen T, et al. Colloidal magnetic resonance contrast agents: effect of particle surface on biodistribution. J Magn Magn Mater 1993, 122:383–386.

29. von zur Muhlen C, von Elverfeldt D, Bassler N, Neudorfer I, Steitz B, et al. Superparamagnetic iron oxide binding and uptake as imaged by magnetic resonance is mediated by the integrin receptor Mac-1 (CD11b/CD18): implications on imaging of atherosclerotic plaques. Atherosclerosis 2007, 193(1):102–111.

17. Wilhelm C, Billotey C, Roger J, Pons JN, Bacri JC, et al. Intracellular uptake of anionic superparamagnetic nanoparticles as a function of their surface coating. Biomaterials 2003, 24(6):1001–1011.

30. Metz S, Bonaterra G, Rudelius M, Settles M, Rummeny EJ, et al. Capacity of human monocytes to phagocytose approved iron oxide MR contrast agents in vitro. Eur Radiol 2004, 14(10):1851–1858.

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

291

Advanced Review

www.wiley.com/wires/nanomed

31. Moffat BA, Reddy GR, McConville P, Hall DE, Chenevert TL, et al. A novel polyacrylamide magnetic nanoparticle contrast agent for molecular imaging using MRI. Mol Imaging 2003, 2(4):324–332. 32. Shi X, Thomas TP, Myc LA, Kotlyar A, Baker JR Jr Synthesis, characterization, and intracellular uptake of carboxyl-terminated poly(amidoamine) dendrimerstabilized iron oxide nanoparticles. Phys Chem Chem Phys 2007, 9(42):5712–5720. 33. Shapiro EM, Skrtic S, Sharer K, Hill JM, Dunbar CE, et al. MRI detection of single particles for cellular imaging. Proc Natl Acad Sci U S A 2004, 101((30):):10901–10906. 34. Shapiro EM, Skrtic S, Koretsky AP. Sizing it up: cellular MRI using micron sized iron oxide particles. Magn Reson Med 2005, 53(2):329–338. 35. Wu YL, Ye Q, Foley LM, Hitchens TK, Sato K, et al. In situ labeling of immune cells with iron oxide particles: an approach to detect organ rejection by cellular MRI. Proc Natl Acad Sci U S A 2006, 103(6):1852–1857. 36. Williams JB, Ye Q, Hitchens TK, Kaufman CL, Ho C. MRI detection of macrophages labeled using micrometer-sized iron oxide particles. J Magn Reson Imaging 2007, 25(6):1210–1218. 37. Shapiro EM, Sharer K, Skrtic S, Koretsky AP. In vivo detection of single cells by MRI. Magn Reson Med 2006, 55(2):242–249. 38. Shapiro EM, Medford-Davis LN, Fahmy TM, Dunbar CE, Koretsky AP. Antibody-mediated cell labeling of peripheral T cells with micron-sized iron oxide particles (MPIOs) allows single cell detection by MRI. Contrast Media Mol Imaging 2007, 2(3):147–153. 39. Slotkin JR, Cahill KS, Tharin SA, Shapiro EM. Cellular magnetic resonance imaging: nanometer and micrometer size particles for noninvasive cell localization. Neurotherapeutics 2007, 4(3):428–433. 40. Jaffer FA, Nahrendorf M, Sosnovik D, Kelly KA, Aikawa E, et al. Cellular imaging of inflammation in atherosclerosis using magnetofluorescent nanomaterials. Mol Imaging 2006, 5(2):85–92. 41. Aikawa E, Nahrendorf M, Sosnovik D, Lok VM, Jaffer FA, et al. Multimodality molecular imaging identifies proteolytic and osteogenic activities in early aortic valve disease. Circulation 2007, 115(3):377–386. 42. Sosnovik DE, Nahrendorf M, Deliolanis N, Novikov M, Aikawa E, et al. Fluorescence tomography and magnetic resonance imaging of myocardial macrophage infiltration in infarcted myocardium in vivo. Circulation 2007, 115(11):1384–1391. 43. Kircher MF, Josephson L, Weissleder R. Ratio imaging of enzyme activity using dual wavelength optical reporters. Mol Imaging 2002, 1(2):89–95.

292

44. Misselwitz B, Platzek J, Weinmann HJ. Early MR lymphography with gadofluorine M in rabbits. Radiology 2004, 231(3):682–688. 45. Henning TD, Saborowski O, Golovko D, Boddington S, Bauer JS, et al. Cell labeling with the positive MR contrast agent gadofluorine M. Eur Radiol 2007, 17(5):1226–1234. 46. Lisy MR, Hartung A, Lang C, Schuler D, Richter W, et al. Fluorescent bacterial magnetic nanoparticles as bimodal contrast agents. Invest Radiol 2007, 42(4):235–241. 47. Nakashima-Matsushita N, Homma T, Yu S, Matsuda T, Sunahara N, et al. Selective expression of folate receptor beta and its possible role in methotrexate transport in synovial macrophages from patients with rheumatoid arthritis. Arthritis Rheum 1999, 42(8):1609–1616. 48. Antony AC. Folate receptors. Annu Rev Nutr 1996, 16:501–521. 49. Saborowski O, Simon GH, Raatschen HJ, Wendland MF, Fu Y, et al. MR imaging of antigen-induced arthritis with a new, folate receptor-targeted contrast agent. Contrast Media Mol Imaging 2007, 2(2):72–81. 50. Ross JF, Chaudhuri PK, Ratnam M. Differential regulation of folate receptor isoforms in normal and malignant tissues in vivo and in established cell lines. Physiologic and clinical implications. Cancer 1994, 73(9):2432–2443. 51. Gustafsson B, Youens S, Louie AY. Development of contrast agents targeted to macrophage scavenger receptors for MRI of vascular inflammation. Bioconjug Chem 2006, 17(2):538–547. 52. Haberland ME, Tannenbaum CS, Williams RE, Adams DO, Hamilton TA. Role of the maleylalbumin receptor in activation of murine peritoneal macrophages in vitro. J Immunol 1989, 142(3):855–862. 53. Ottnad E, Via DP, Frubis J, Sinn H, Friedrich E, et al. Differentiation of binding sites on reconstituted hepatic scavenger receptors using oxidized low-density lipoprotein. Biochem J 1992, 281(3):745–751. 54. Briley-Saebo KC, Amirbekian V, Mani V, Aguinaldo JG, Vucic E, et al. Gadolinium mixed-micelles: effect of the amphiphile on in vitro and in vivo efficacy in apolipoprotein E knockout mouse models of atherosclerosis. Magn Reson Med 2006, 56(6):1336–1346. 55. Lipinski MJ, Amirbekian V, Frias JC, Aguinaldo JG, Mani V, et al. MRI to detect atherosclerosis with gadolinium-containing immunomicelles targeting the macrophage scavenger receptor. Magn Reson Med 2006, 56(3):601–610. 56. Amirbekian V, Lipinski MJ, Briley-Saebo KC, Amirbekian S, Aguinaldo JG, et al. Detecting and assessing

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

macrophages in vivo to evaluate atherosclerosis noninvasively using molecular MRI. Proc Natl Acad Sci U S A 2007, 104(3):961–966. 57. Cunningham CH, Arai T, Yang PC, McConnell MV, Pauly JM, et al. Positive contrast magnetic resonance imaging of cells labeled with magnetic nanoparticles. Magn Reson Med 2005, 53(5):999–1005. 58. Heyn C, Bowen CV, Rutt BK, Foster PJ. Detection threshold of single SPIO-labeled cells with FIESTA. Magn Reson Med 2005, 53(2):312–320. ¨ M, Kedziorek DA, Hof59. Stuber M, Gilson WD, Schar mann LV, et al. Positive contrast visualization of iron oxide-labeled stem cells using inversion-recovery with ON-resonant water suppression (IRON). Magn Reson Med 2007, 58(5):1072–1077. 60. Seppenwoolde JH, Viergever MA, Bakker CJ. Passive tracking exploiting local signal conservation: the white marker phenomenon. Magn Reson Med 2003, 50(4):784–790. 61. Mani V, Briley-Saebo KC, Itskovich VV, Samber DD, Fayad ZA. Gradient echo acquisition for superparamagnetic particles with positive contrast (GRASP): sequence characterization in membrane and glass superparamagnetic iron oxide phantoms at 1.5T and 3T. Magn Reson Med 2006, 55(1):126–135. 62. Weissleder R, Elizondo G, Wittenberg J, Rabito CA, Bengele HH, et al. Ultrasmall superparamagnetic iron oxide: characterization of a new class of contrast agents for MR imaging. Radiology 1990, 175(2):489–493. 63. Dousset V, Gomez C, Petry KG, Delalande C, Caille JM. Dose and scanning delay using USPIO for central nervous system macrophage imaging. Magma 1999, 8(3):185–189. ¨ 64. Oude Engberink RD, van der Pol SM, Dopp EA, de Vries HE, Blezer EL. Comparison of SPIO and USPIO for in vitro labeling of human monocytes: MR detection and cell function. Radiology 2007, 243(2):467–474. 65. Schroeter M, Saleh A, Wiedermann D, Hoehn M, Jander S. Histochemical detection of ultrasmall superparamagnetic iron oxide (USPIO) contrast medium uptake in experimental brain ischemia. Magn Reson Med 2004, 52(2):403–406. 66. Hansson GK, Libby P, Schonbeck U, Yan ZQ. Innate and adaptive immunity in the pathogenesis of atherosclerosis. Circ Res 2002, 91(4):281–291.

atherogenesis and the metabolism of modified LDL and HDL. Ann N Y Acad Sci 2000, 902:113–126. 70. Kooi ME, Cappendijk VC, Cleutjens KB, Kessels AG, Kitslaar PJ, et al. Accumulation of ultrasmall superparamagnetic particles of iron oxide in human atherosclerotic plaques can be detected by in vivo magnetic resonance imaging. Circulation 2003, 107(19):2453–2458. 71. Corot C, Petry KG, Trivedi R, Saleh A, Jonkmanns C, et al. Macrophage imaging in central nervous system and in carotid atherosclerotic plaque using ultrasmall superparamagnetic iron oxide in magnetic resonance imaging. Invest Radiol 2004, 39(10):619–625. 72. Ruehm SG, Corot C, Vogt P, Kolb S, Debatin JF. Magnetic resonance imaging of atherosclerotic plaque with ultrasmall superparamagnetic particles of iron oxide in hyperlipidemic rabbits. Circulation 2001, 103(3):415–422. 73. Rogers WJ, Basu P. Factors regulating macrophage endocytosis of nanoparticles: implications for targeted magnetic resonance plaque imaging. Atherosclerosis 2005, 178(1):67–73. 74. Anzai Y, Prince MR, Chenevert TL, Maki JH, Londy F, et al. MR angiography with an ultrasmall superparamagnetic iron oxide blood pool agent. J Magn Reson Imaging 1997, 7(1):209–214. 75. Hyafil F, Cornily JC, Feig JE, Gordon R, Vucic E, et al. Noninvasive detection of macrophages using a nanoparticulate contrast agent for computed tomography. Nat Med 2007, 13(5):636–641. 76. Schmitz SA, Coupland SE, Gust R, Winterhalter S, Wagner S, et al. Superparamagnetic iron oxideenhanced MRI of atherosclerotic plaques in Watanabe hereditable hyperlipidemic rabbits. Invest Radiol 2000, 35(8):460–471. 77. Yancy AD, Olzinski AR, Hu TC, Lenhard SC, Aravindhan K, et al. Differential uptake of ferumoxtran-10 and ferumoxytol, ultrasmall superparamagnetic iron oxide contrast agents in rabbit: critical determinants of atherosclerotic plaque labeling. J Magn Reson Imaging 2005, 21(4):432–442. 78. Herborn CU, Vogt FM, Lauenstein TC, Dirsch O, Corot C, et al. Magnetic resonance imaging of experimental atherosclerotic plaque: comparison of two ultrasmall superparamagnetic particles of iron oxide. J Magn Reson Imaging 2006, 24(2):388–393.

67. Stoll G, Bendszus M. Inflammation and atherosclerosis: novel insights into plaque formation and destabilization. Stroke 2006, 37(7):1923–1932.

79. Munoz LE, Frey B, Pausch F, Baum W, Mueller RB, et al. The role of annexin A5 in the modulation of the immune response against dying and dead cells. Curr Med Chem 2007, 14(3):271–277.

68. Virmani R, Burke AP, Farb A, Kolodgie FD. Pathology of the vulnerable plaque. J Am Coll Cardiol 2006, 47(suppl 8):C13–C18.

80. Laufer EM, Reutelingsperger CP, Narula J, Hofstra L. Annexin A5: an imaging biomarker of cardiovascular risk. Basic Res Cardiol 2008, 103(2):95–104.

69. van Berkel TJ, van Eck M, Herijgers N, Fluiter K, Nion S. Scavenger receptor classes A and B. Their roles in

81. Corsten MF, Reutelingsperger CP, Hofstra L. Imaging apoptosis for detecting plaque instability: rendering

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

293

Advanced Review

www.wiley.com/wires/nanomed

death a brighter facade. Curr Opin Biotechnol 2007, 18(1):83–89. 82. Smith BR, Heverhagen J, Knopp M, Schmalbrock P, Shapiro J, et al. Localization to atherosclerotic plaque and biodistribution of biochemically derivatized superparamagnetic iron oxide nanoparticles (SPIONs) contrast particles for magnetic resonance imaging (MRI). Biomed Microdevices 2007, 9(5):719–727. 83. Sosnovik DE, Schellenberger EA, Nahrendorf M, Novikov MS, Matsui T, et al. Magnetic resonance imaging of cardiomyocyte apoptosis with a novel magneto-optical nanoparticle. Magn Reson Med 2005, 54(3):718–724.

93. Aktas O, Ullrich O, Infante-Duarte C, Nitsch R, Zipp F. Neuronal damage in brain inflammation. Arch Neurol 2007, 64(2):185–189. 94. Muir KW, Tyrrell P, Sattar N, Warburton E. Inflammation and ischaemic stroke. Curr Opin Neurol 2007, 20(3):334–342. 95. Wang Q, Tang XN, Yenari MA. The inflammatory response in stroke. J Neuroimmunol 2007, 184(1–2):53–68. 96. Lopez-Neblina F, Toledo AH, Toledo-Pereyra LH. Molecular biology of apoptosis in ischemia and reperfusion. J Invest Surg 2005, 18(6):335–350.

84. Schellenberger EA, Sosnovik D, Weissleder R, Josephson L. Magneto/optical annexin V, a multimodal protein. Bioconjug Chem 2004, 15(5):1062–1067.

97. Bazan NG, Marcheselli VL, Cole-Edwards K. Brain response to injury and neurodegeneration: endogenous neuroprotective signaling. Ann N Y Acad Sci 2005, 1053:137–147.

85. Trivedi RA, U-King-Im JM, Graves MJ, Kirkpatrick PJ, Gillard JH. Noninvasive imaging of carotid plaque inflammation. Neurology 2004, 63(1):187–188.

98. Zheng Z, Yenari MA. Post-ischemic inflammation: molecular mechanisms and therapeutic implications. Neurol Res 2004, 26(8):884–892.

86. Trivedi RA, U-King-Im JM, Graves MJ, Cross JJ, Horsley J, et al. In vivo detection of macrophages in human carotid atheroma: temporal dependence of ultrasmall superparamagnetic particles of iron oxideenhanced MRI. Stroke 2004, 35(7):1631–1635.

99. Stoll G. Inflammatory cytokines in the nervous system: multifunctional mediators in autoimmunity and cerebral ischemia. Rev Neurol (Paris) 2002, 158(10 Pt 1):887–891.

87. Trivedi RA, Mallawarachi C, U-King-Im JM, Graves MJ, Horsley J, et al. Identifying inflamed carotid plaques using in vivo USPIO-enhanced MR imaging to label plaque macrophages. Arterioscler Thromb Vasc Biol 2006, 26(7):1601–1606. 88. Tang T, Howarth SP, Miller SR, Trivedi R, Graves MJ, et al. Assessment of inflammatory burden contralateral to the symptomatic carotid stenosis using highresolution ultrasmall, superparamagnetic iron oxideenhanced MRI. Stroke 2006, 37(9):2266–2270. 89. Howarth S, Li ZY, Trivedi RA, U-King-Im JM, Graves MJ, et al. Correlation of macrophage location and plaque stress distribution using USPIO-enhanced MRI in a patient with symptomatic severe carotid stenosis: a new insight into risk stratification. Br J Neurosurg 2007, 21(4):396–398. 90. ClinicalTrials.gov. Effects of atorvastatin on macrophage activity and plaque inflammation using magnetic resonance imaging. http://www. clinicaltrials.gov. Accessed November 2, 2007. 91. Kerwin W, Hooker A, Spilker M, Vicini P, Ferguson M, et al. Quantitative magnetic resonance imaging analysis of neovasculature volume in carotid atherosclerotic plaque. Circulation 2003, 107(6):851–856. 92. Kerwin WS, O’Brien KD, Ferguson MS, Polissar N, Hatsukami TS, et al. Inflammation in carotid atherosclerotic plaque: a dynamic contrastenhanced MR imaging study. Radiology 2006, 241(2):459–468.

294

100. Schilling M, Besselmann M, Leonhard C, Mueller M, Ringelstein EB, et al. Microglial activation precedes and predominates over macrophage infiltration in transient focal cerebral ischemia: a study in green fluorescent protein transgenic bone marrow chimeric mice. Exp Neurol 2003, 183(1):25–33. 101. Mehta SL, Manhas N, Raghubir R. Molecular targets in cerebral ischemia for developing novel therapeutics. Brain Res Rev 2007, 54(1):34–66. 102. Kleinschnitz C, Bendszus M, Frank M, Solymosi L, Toyka KV, et al. In vivo monitoring of macrophage infiltration in experimental ischemic brain lesions by magnetic resonance imaging. J Cereb Blood Flow Metab 2003, 23(11):1356–1361. 103. Saleh A, Wiedermann D, Schroeter M, Jonkmanns C, Jander S, et al. Central nervous system inflammatory response after cerebral infarction as detected by magnetic resonance imaging. NMR Biomed 2004, 17(4):163–169. 104. Saleh A, Schroeter M, Ringelstein A, Hartung HP, Siebler M, et al. Iron oxide particle-enhanced MRI suggests variability of brain inflammation at early stages after ischemic stroke. Stroke 2007, 38(10):2733–2737. 105. Jander S, Schroeter M, Saleh A. Imaging inflammation in acute brain ischemia. Stroke 2007, 38(suppl 2):642–645. 106. Wiart M, Davoust N, Pialat JB, Desestret V, Moucharaffie S, et al. MRI monitoring of neuroinflammation in mouse focal ischemia. Stroke 2007, 38(1):131–137.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

107. Kleinschnitz C, Schutz A, Nolte I, Horn T, Frank M, et al. In vivo detection of developing vessel occlusion in photothrombotic ischemic brain lesions in the rat by iron particle enhanced MRI. J Cereb Blood Flow Metab 2005, 25(11):1548–1555. ¨ 108. Rausch M, Sauter A, Frohlich J, Neubacher U, Radu¨ EW, et al. Dynamic patterns of USPIO enhancement can be observed in macrophages after ischemic brain damage. Magn Reson Med 2001, 46(5):1018–1022.

119. Hendriks JJ, Teunissen CE, de Vries HE, Dijkstra CD. Macrophages and neurodegeneration. Brain Res Brain Res Rev 2005, 48(2):185–195. 120. Barnett MH, Henderson AP, Prineas JW. The macrophage in MS: just a scavenger after all? Pathology and pathogenesis of the acute MS lesion. Mult Scler 2006, 12(2):121–132. 121. Lassmann H. Experimental models of multiple sclerosis. Rev Neurol (Paris) 2007, 163(6–7):651–655.

109. Rausch M, Baumann D, Neubacher U, Rudin M. Invivo visualization of phagocytotic cells in rat brains after transient ischemia by USPIO. NMR Biomed 2002, 15(4):278–283.

122. Dousset V, Delalande C, Ballarino L, Quesson B, Seilhan D, et al. In vivo macrophage activity imaging in the central nervous system detected by magnetic resonance. Magn Reson Med 1999, 41(2):329–333.

110. Neumann-Haefelin T, Kastrup A, de Crespigny A, Yenari MA, Ringer T, et al. Serial MRI after transient focal cerebral ischemia in rats: dynamics of tissue injury, blood-brain barrier damage, and edema formation. Stroke 2000, 31(8):1965–1972.

123. Dousset V, Ballarino L, Delalande C, Coussemacq M, Canioni P, et al. Comparison of ultrasmall particles of iron oxide (USPIO) T2-weighted, conventional T2-weighted and gadolinium-enhanced T1-weighted MR images in rats with experimental autoimmune encephalomyelitis. Am J Neuroradiol 1999, 20(2):223–227.

111. Engberink RD, Blezer EL, Hoff EI, van der Pol SM, van der Toorn A, et al. MRI of monocyte infiltration in an animal model of neuroinflammation using SPIOlabeled monocytes or free USPIO. J Cereb Blood Flow Metab 2008, 28(4):841–851. 112. Schroeter M, Jander S, Huitinga I, Witte OW, Stoll G. Phagocytic response in photochemically induced infarction of the rat cerebral cortex: the role of resident microglia. Stroke 1997, 28(2):382–386. 113. Tanaka R, Komine-Kobayashi M, Mochizuki H, Yamada M, Furuya T, et al. Migration of enhanced green fluorescent protein expressing bone marrowderived microglia/macrophage into the mouse brain following permanent focal ischemia. Neuroscience 2003, 117(3):531–539. 114. Saleh A, Schroeter M, Jonkmanns C, Hartung HP, Modder U, et al. In vivo MRI of brain inflammation in human ischaemic stroke. Brain 2004, 127(7):1670–1677. 115. Nighoghossian N, Wiart M, Cakmak S, Berthezene Y, Derex L, et al. Inflammatory response after ischemic stroke: a USPIO-enhanced MRI study in patients. Stroke 2007, 38(2):303–307. 116. Beaulieu C, de Crespigny A, Tong DC, Moseley ME, Albers GW, et al. Longitudinal magnetic resonance imaging study of perfusion and diffusion in stroke: evolution of lesion volume and correlation with clinical outcome. Ann Neurol 1999, 46(4):568–578. 117. Dousset V, Brochet B, Deloire MS, Lagoarde L, Barroso B, et al. MR imaging of relapsing multiple sclerosis patients using ultra-small-particle iron oxide and compared with gadolinium. Am J Neuroradiol 2006, 27(5):1000–1005. 118. Raivich G, Banati R. Brain microglia and bloodderived macrophages: molecular profiles and functional roles in multiple sclerosis and animal models of autoimmune demyelinating disease. Brain Res Brain Res Rev 2004, 46(3):261–281.

Vo lu me 1, May /Ju n e 2009

124. Rausch M, Hiestand P, Baumann D, Cannet C, Rudin M. MRI-based monitoring of inflammation and tissue damage in acute and chronic relapsing EAE. Magn Reson Med 2003, 50(2):309–314. 125. Oweida AJ, Dunn EA, Karlik SJ, Dekaban GA, Foster PJ. Iron-oxide labeling of hematogenous macrophages in a model of experimental autoimmune encephalomyelitis and the contribution to signal loss in fast imaging employing steady state acquisition (FIESTA) images. J Magn Reson Imaging 2007, 26(1):144–151. ¨ 126. Linker RA, Kroner A, Horn T, Gold R, Maurer M, et al. Iron particle-enhanced visualization of inflammatory central nervous system lesions by high resolution sonography: preliminary data in an animal model. Am J Neuroradiol 2006, 27(6):1225–1229. 127. Berger C, Hiestand P, Kindler-Baumann D, Rudin M, Rausch M. Analysis of lesion development during acute inflammation and remission in a rat model of experimental autoimmune encephalomyelitis by visualization of macrophage infiltration, demyelination and blood-brain barrier damage. NMR Biomed 2006, 19(1):101–107. 128. de Vries HE, Kuiper J, de Boer AG, Van Berkel TJ, Breimer DD. The blood-brain barrier in neuroinflammatory diseases. Pharmacol Rev 1997, 49(2):143–155. 129. Floris S, Blezer EL, Schreibelt G, Dopp E, van der Pol SM, et al. Blood-brain barrier permeability and monocyte infiltration in experimental allergic encephalomyelitis: a quantitative MRI study. Brain 2004, 127(3):616–627. 130. Brochet B, Deloire MS, Touil T, Anne O, Caille JM, et al. Early macrophage MRI of inflammatory lesions predicts lesion severity and disease development in relapsing EAE. Neuroimage 2006, 32(1):266–274.

 2009 Jo h n Wiley & So n s, In c.

295

Advanced Review

www.wiley.com/wires/nanomed

131. Rausch M, Hiestand P, Foster CA, Baumann D, Cannet C, et al. Predicability of FTY720 efficacy in experimental autoimmune encephalomyelitis by in vivo macrophage tracking: clinical implications for ultrasmall superparamagnetic iron oxide-enhanced magnetic resonance imaging. J Magn Reson Imaging 2004, 20(1):16–24. 132. Deloire MS, Touil T, Brochet B, Dousset V, Caill´e JM, et al. Macrophage brain infiltration in experimental autoimmune encephalomyelitis is not completely compromised by suppressed T-cell invasion: in vivo magnetic resonance imaging illustration in effective anti-VLA-4 antibody treatment. Mult Scler 2004, 10(5):540–548. 133. Petry KG, Boiziau C, Dousset V, Brochet B. Magnetic resonance imaging of human brain macrophage infiltration. Neurotherapeutics 2007, 4(3):434–442. 134. Manninger SP, Muldoon LL, Nesbit G, Murillo T, Jacobs PM, et al. An exploratory study of ferumoxtran-10 nanoparticles as a blood-brain barrier imaging agent targeting phagocytic cells in CNS inflammatory lesions. Am J Neuroradiol 2005, 26(9):2290–2300. 135. Vellinga MM, Oude Engberink RD, Seewann A, Pouwels PJ, Wattjes MP, et al. Pluriformity of inflammation in multiple sclerosis shown by ultrasmall iron oxide particle enhancement. Brain 2008, 131(3):800–807. 136. Lucchinetti C, Bruck W, Parisi J, Scheithauer B, Rodriguez M, et al. Heterogeneity of multiple sclerosis lesions; implications for the pathogenesis of demyelination. Ann Neurol 2000, 47(6):707–717. 137. Morales Y, Parisi JE, Lucchinetti CF. The pathology of multiple sclerosis: evidence for heterogeneity. Adv Neurol 2006, 98:27–45. 138. Hohlfeld R. Does inflammation stimulate remyelination? J Neurol 2007, 254(suppl 1):I47–I54. 139. Nessler S, Boretius S, Stadelmann C, Bittner A, Merkler D, et al. Early MRI changes in a mouse model of multiple sclerosis are predictive of severe inflammatory tissue damage. Brain 2007, 130(8):2186–2198. 140. Silver NC, Good CD, Sormani MP, MacManus DG, Thompson AJ, et al. A modified protocol to improve the detection of enhancing brain and spinal cord lesions in multiple sclerosis. J Neurol 2001, 248(3):215–224. 141. Vos CM, Geurts JJ, Montagne L, van Haastert ES, Bo L, et al. Blood-brain barrier alterations in both focal and diffuse abnormalities on postmortem MRI in multiple sclerosis. Neurobiol Dis 2005, 20(3):953–960. 142. Cutolo M. Macrophages as effectors of the immunoendocrinologic interactions in autoimmune rheumatic diseases. Ann N Y Acad Sci 1999, 876:32–41. 143. Tak PP, Smeets TJ, Daha MR, Kluin PM, Meijers KA, et al. Analysis of the synovial cell infiltrate in early

296

rheumatoid synovial tissue in relation to local disease activity. Arthritis Rheum 1997, 40(2):217–225. 144. Kinne RW, Brauer R, Stuhlmuller B, Palombo-Kinne E, Burmester GR. Macrophages in rheumatoid arthritis. Arthritis Res 2000, 2(3):189–202. 145. Dardzinski BJ, Schmithorst VJ, Holland SK, Boivin GP, Imagawa T, et al. MR imaging of murine arthritis using ultrasmall superparamagnetic iron oxide particles. Magn Reson Imaging 2001, 19(9):1209–1216. 146. Beckmann N, Kneuer R, Gremlich HU, KarmoutyQuintana H, Bl´e FX, et al. In vivo mouse imaging and spectroscopy in drug discovery. NMR Biomed 2007, 20(3):154–185. 147. Beckmann N, Falk R, Zurbrugg S, Dawson J, Engelhardt P. Macrophage infiltration into the rat knee detected by MRI in a model of antigen-induced arthritis. Magn Reson Med 2003, 49(6):1047–1055. 148. Lutz AM, Seemayer C, Corot C, Gay RE, Goepfert K, et al. Detection of synovial macrophages in an experimental rabbit model of antigen-induced arthritis: ultrasmall superparamagnetic iron oxide-enhanced MR imaging. Radiology 2004, 233(1):149–157. 149. Wooley PH. The usefulness and the limitations of animal models in identifying targets for therapy in arthritis. Best Pract Res Clin Rheumatol 2004, 18(1):47–58. 150. Simon GH, von Vopelius-Feldt J, Wendland MF, Fu Y, Piontek G, et al. MRI of arthritis: comparison of ultrasmall superparamagnetic iron oxide vs. GdDTPA. J Magn Reson Imaging 2006, 23(5):720–727. 151. Hansch A, Frey O, Sauner D, Hilger I, Haas M: et al. In vivo imaging of experimental arthritis with near-infrared fluorescence. Arthritis Rheum 2004, 50(3):961–967. 152. Chen WT, Mahmood U, Weissleder R, Tung CH. Arthritis imaging using a near infrared fluorescence folate-targeted probe. Arthritis Res Ther 2005, 7(2):R310–R317. 153. Paulos CM, Turk MJ, Breur GJ, Low PS. Folate receptor-mediated targeting of therapeutic and imaging agents to activated macrophages in rheumatoid arthritis. Adv Drug Deliv Rev 2004, 56(8):1205–1217. 154. Quinn MA, Emery P. Window of opportunity in early rheumatoid arthritis: possibility of altering the disease process with early intervention. Clin Exp Rheumatol 2003, 21(5, suppl 31):S154–S157. 155. Lucas T, Abraham D, Aharinejad S. Modulation of tumor associated macrophages in solid tumors. Front Biosci 2008, 13:5580–5588. 156. Noonan DM, De Lerma Barbaro A, Vannini N, Mortara L, Albini A. Inflammation, inflammatory cells and angiogenesis: decisions and indecisions. Cancer Metastasis Rev 2008, 27(1):31–40.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

In Vivo Visualization of Macrophage Infiltration and Activity

WIREs Nanomedicine and Nanobiotechnology

157. Mantovani A, Marchesi F, Porta C, Sica A, Allavena P. Inflammation and cancer: breast cancer as a prototype. Breast 2007, 16(suppl 2):S27–S33. 158. Wagenlehner FM, Elkahwaji JE, Algaba F, BjerklundJohansen T, Naber KG, et al. The role of inflammation and infection in the pathogenesis of prostate carcinoma. BJU Int 2007, 100(4):733–737. 159. Kremer S, Pinel S, V´edrine PO, Bressenot A, Robert P, et al. Ferumoxtran-10 enhancement in orthotopic xenograft models of human brain tumors: an indirect marker of tumor proliferation? J Neurooncol 2007, 83(2):111–119. 160. Valable S, Barbier EL, Bernaudin M, Roussel S, Segebarth C, et al. In vivo MRI tracking of exogenous monocytes/macrophages targeting brain tumors in a rat model of glioma. Neuroimage 2008, 40(2):973–983. 161. Guimaraes R, Cl´ement O, Bittoun J, Carnot F, Frija G. MR lymphography with superparamagnetic iron nanoparticles in rats: pathologic basis for contrast enhancement. Am J Roentgenol 1994, 162(1):201–207. 162. Vassallo P, Matei C, Heston WD, McLachlan SJ, Koutcher JA, et al. AMI-227-enhanced MR lymphography: usefulness for differentiating reactive from tumor-bearing lymph nodes. Radiology 1994, 193(2):501–506. ¨ 163. Barentsz JO, Futterer JJ, Takahashi S. Use of ultrasmall superparamagnetic iron oxide in lymph node MR imaging in prostate cancer patients. Eur J Radiol 2007, 63(3):369–372. 164. Bellin MF, Lebleu L, Meric JB. Evaluation of retroperitoneal and pelvic lymph node metastases with MRI and MR lymphangiography. Abdom Imaging 2003, 28(2):155–163. 165. Tatsumi Y, Tanigawa N, Nishimura H, Nomura E, Mabuchi H, et al. Preoperative diagnosis of lymph node metastases in gastric cancer by magnetic resonance imaging with ferumoxtran-10. Gastric Cancer 2006, 9(2):120–128. 166. Nishimura H, Tanigawa N, Hiramatsu M, Tatsumi Y, Matsuki M, et al. Preoperative esophageal cancer staging: magnetic resonance imaging of lymph node with ferumoxtran-10, an ultrasmall superparamagnetic iron oxide. J Am Coll Surg 2006, 202(4):604–611. ´ 167. Neuwelt EA, Varallyay P, Bago´ AG, Muldoon LL, Nesbit G, et al. Imaging of iron oxide nanoparticles by MR and light microscopy in patients with malignant brain tumours. Neuropathol Appl Neurobiol 2004, 30(5):456–471. 168. Taschner CA, Wetzel SG, Tolnay M, Froehlich J, Merlo A, et al. Characteristics of ultrasmall superparamagnetic iron oxides in patients with brain tumors. Am J Roentgenol 2005, 185(6):1477–1486.

Vo lu me 1, May /Ju n e 2009

169. Schuurman HJ, Weckbecker G, Bruns C, Beckmann N, Rudin M, et al. Preclinical models of chronic rejection: promises and pitfalls. Transplant Proc 1997, 29(6):2624–2625. 170. Wu YJ, Sato K, Ye Q, Ho C. MRI investigations of graft rejection following organ transplantation using rodent models. Methods Enzymol 2004, 386:73–105. 171. Rocha PN, Plumb TJ, Crowley SD, Coffman TM. Effector mechanisms in transplant rejection. Immunol Rev 2003, 196:51–64. 172. Grau V, Herbst B, Steiniger B. Dynamics of monocytes/macrophages and T lymphocytes in acutely rejecting rat renal allografts. Cell Tissue Res 1998, 291(1):117–126. 173. Croker BP, Clapp WL, Abu Shamat AR, Kone BC, Peterson JC. Macrophages and chronic renal allograft nephropathy. Kidney Int Suppl 1996, 57:S42–S49. 174. Azuma H, Takahara S, Matsumoto K, Ichimaru N, Wang JD, et al. Hepatocyte growth factor prevents the development of chronic allograft nephropathy in rats. J Am Soc Nephrol 2001, 12(6):1280–1292. 175. Herrero-Fresneda I, Torras J, Cruzado JM, Condom E, Vidal A, et al. Do alloreactivity and prolonged cold ischemia cause different elementary lesions in chronic allograft nephropathy? Am J Pathol 2003, 162(1):127–137. ¨ 176. Hamar P, Szabo A, Muller V, Heemann U. Involvement of interleukin-2 and growth factors in chronic kidney allograft rejection in rats. Transplant Proc 2001, 33(3):2160–2162. 177. Ziai F, Nagano H, Kusaka M, Coito AJ, Troy JL, et al. Renal allograft protection with losartan in Fisher–> Lewis rats: hemodynamics, macrophages, and cytokines. Kidney Int 2000, 57(6):2618–2625. 178. Yang J, Reutzel-Selke A, Steier C, Jurisch A, Tullius SG, et al. Targeting of macrophage activity by adenovirus-mediated intragraft overexpression of TNFRp55-Ig, IL-12p40, and vIL-10 ameliorates adenovirus-mediated chronic graft injury, whereas stimulation of macrophages by overexpression of IFN-gamma accelerates chronic graft injury in a rat renal allograft model. J Am Soc Nephrol 2003, 14(1):214–225. 179. Engelbrecht G, Kahn D, Duminy F, Hickman R. New rapid technique for renal-transplantation in the rat. Microsurgery 1992, 13(6):340–344. 180. Zhang Y, Dodd SJ, Hendrich KS, Williams M, Ho C. Magnetic resonance imaging detection of rat renal transplant rejection by monitoring macrophage infiltration. Kidney Int 2000, 58(3):1300–1310. 181. Beckmann N, Cannet C, Fringeli-Tanner M, Baumann D, Pally C, et al. Macrophage labeling by SPIO as an early marker of allograft chronic rejection in a rat model of kidney transplantation. Magn Reson Med 2003, 49(3):459–467.

 2009 Jo h n Wiley & So n s, In c.

297

Advanced Review

www.wiley.com/wires/nanomed

¨ R, Li 182. Beckmann N, Cannet C, Zurbruegg S, Haberthur J, et al. Macrophage infiltration detected at MR imaging in rat kidney allografts: early marker of chronic rejection? Radiology 2006, 240(3):717–724. 183. Hauger O, Grenier N, Deminere C, Lasseur C, Delmas Y, et al. USPIO-enhanced MR imaging of macrophage infiltration in native and transplanted kidneys: initial results in humans. Eur Radiol 2007, 17(11):2898–2907. 184. Rudin M, Beckmann N, Rausch M. Magnetic resonance imaging in biomedical research: imaging of drugs and drug effects. Methods Enzymol 2004, 385:240–256. 185. Beckmann N, Laurent D, Tigani B, Panizzutti R, Rudin M. Magnetic resonance imaging in drug discovery: lessons from disease areas. Drug Discov Today 2004, 9(1):35–42. 186. Beckmann N: ed.In Vivo MR Techniques in Drug Discovery and Development:New York: Taylor & Francis; 2006. 187. Anzai Y, Piccoli CW, Outwater EK, Stanford W, Bluemke DA, et al. Evaluation of neck and body metastases to nodes with ferumoxtran 10-enhanced MR imaging: phase III safety and efficacy study. Radiology 2003, 228(3):777–788. 188. Harisinghani MG, Barentsz J, Hahn PF, Deserno WM, Tabatabaei S, et al. Noninvasive detection of clinically occult lymph-node metastases in prostate cancer. N Engl J Med 2003, 348(25):2491–2499. 189. Medina C, Santos-Martinez MJ, Radomski A, Corrigan OI, Radomski MW. Nanoparticles: pharmacological and toxicological significance. Br J Pharmacol 2007, 150(5):552–558. ¨ 190. Muller K, Skepper JN, Posfai M, Trivedi R, Howarth S, et al. Effect of ultrasmall superparamagnetic iron oxide nanoparticles (Ferumoxtran-10) on human monocyte-macrophages in vitro. Biomaterials 2007, 28(9):1629–1642. 191. Siglienti I, Bendszus M, Kleinschnitz C, Stoll G. Cytokine profile of iron-laden macrophages: implications for cellular magnetic resonance imaging. J Neuroimmunol 2006, 173(1–2):166–173. 192. Gorantla S, Dou H, Boska M, Destache CJ, Nelson J, et al. Quantitative magnetic resonance and SPECT imaging for macrophage tissue migration and nanoformulated drug delivery. J Leukoc Biol 2006, 80(5):1165–1174.

298

193. Fujita M, Imaizumi M, Zoghbi SS, Fujimura Y, Farris AG, et al. Kinetic analysis in healthy humans of a novel positron emission tomography radioligand to image the peripheral benzodiazepine receptor, a potential biomarker for inflammation. Neuroimage 2008, 40(1):43–52. ¨ 194. Muller C, Forrer F, Schibli R, Krenning EP, de Jong M. SPECT study of folate receptor-positive malignant and normal tissues in mice using a novel 99mTcradiofolate. J Nucl Med 2008, 49(2):310–317. 195. Nahrendorf M, Zhang H, Hembrador S, Panizzi P, Sosnovik DE, et al. Nanoparticle PET-CT imaging of macrophages in inflammatory atherosclerosis. Circulation 2008, 117:379–387. 196. Venneti S, Wang G, Wiley CA. The high affinity peripheral benzodiazepine receptor ligand DAA1106 binds to activated and infected brain macrophages in areas of synaptic degeneration: implications for PET imaging of neuroinflammation in lentiviral encephalitis. Neurobiol Dis 2008, 29(2):232–241. 197. Venneti S, Bonneh-Barkay D, Lopresti BJ, Bissel SJ, Wang G, et al. Longitudinal in vivo positron emission tomography imaging of infected and activated brain macrophages in a macaque model of human immunodeficiency virus encephalitis correlates with central and peripheral markers of encephalitis and areas of synaptic degeneration. Am J Pathol 2008, 172(6):1603–1616. 198. Myers R, Manjil LG, Cullen BM, Price GW, Frackowiak RS, et al. Macrophage and astrocyte populations in relation to 3H-PK 11195 binding in rat cerebral cortex following a local ischaemic lesion. J Cereb Blood Flow Metab 1991, 11:314–322. 199. Stephenson DT, Schober DA, Smalstig EB, Mincy RE, Gehlert DR, et al. Peripheral benzodiazepine receptors are colocalized with activated microglia following transient global forebrain ischemia in the rat. J Neurosci 1995, 15(7 Pt 2):5263–5274. 200. Venneti S, Wang G, Wiley CA. Activated macrophages in HIV encephalitis and a macaque model show increased [3H](R)-PK11195 binding in a PI3kinase-dependent manner. Neurosci Lett 2007, 426(2):117–122. 201. Banati RB. Visualising microglial activation in vivo. Glia 2002, 40(2):206–217. 202. Shashkin P, Dragulev B, Ley K. Macrophage differentiation to foam cells. Curr Pharm Des 2005, 11(23):3061–3072.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

Advanced Review

Magnetic resonance relaxation properties of superparamagnetic particles Yves Gossuin,1∗ Pierre Gillis,1 Aline Hocq,1 Quoc L Vuong1 and Alain Roch2 Nanometric crystals of maghemite are known to exhibit superparamagnetism. Because of the significance of their magnetic moment, maghemite nanoparticles are exceptional contrast agents and are used for magnetic resonance imaging (of the liver, spleen, lymph nodes), for magnetic resonance angiography and for molecular and cellular imaging. The relaxivity of these agents depends on their size, saturation magnetization and magnetic field and also on their degree of clustering. There are different types of maghemite particles whose relaxation characteristics are suited to a specific MRI application. The relaxation induced by maghemite particles is caused by the diffusion of water protons in the inhomogeneous field surrounding the particles. This is well described by a theoretical model that ´ relaxation into account. Another type takes magnetite crystal anisotropy and Neel of superparamagnetic compound is ferritin, the iron-storing protein: it contains a superparamagnetic ferrihydrite core. Even if the resulting magnetic moment of ferritin is far smaller than for magnetite nanoparticles, its massive presence in different organs darkens T2 -weighted MR images, allowing the noninvasive estimation of iron content, thanks to MRI. The relaxation induced by ferritin in aqueous solutions has been demonstrated to be caused by the exchange of protons between bulk water protons and the surface of the ferrihydrite crystal. However, in vivo, the relaxation properties of ferritin are still unexplained, probably because of protein clustering.  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 299–310

F

rom the dawn of Magnetic Resonance Imaging (MRI), magnetic compounds have been used to accelerate proton relaxation.1 Paramagnetic gadolinium complexes are still widely used as positive contrast agents for routine MRI. However, the efficiency of the rare earth ions is limited, mainly because their magnetic moment is not saturated at imaging magnetic fields.2 The second category of contrast agents, on which we will focus in this review, is that of superparamagnetic maghemite nanoparticles—excellent negative agents for MRI.3 These compounds are also the

∗ Correspondence

to: [email protected]

1

Biological Physics Department, University of Mons-Hainaut, Mons, Belgium 2

Department of General, Organic and Biomedical Chemistry, NMR and Molecular Imaging Laboratory, University of Mons-Hainaut, Mons, Belgium DOI: 10.1002/wnan.036

Vo lu me 1, May /Ju n e 2009

best candidates for Magnetic Resonance (MR) cellular and molecular imaging.4 In these applications, the contrast agent is chemically linked to a vector, which will target a specific macromolecule (molecular imaging) or a specific type of cell (cellular imaging). In this way, different biological processes typical of certain diseases can be visualized noninvasively using MRI. Superparamagnetic particles present an important advantage compared to paramagnetic ions since each vectorized particle bears a huge magnetic moment, compared to a single targeted paramagnetic ion. Another example of superparamagnetic particles is found inside endogenous ferritin, whose presence influences the contrast of MR images of liver, spleen and brain.5 This natural contrast has been used to evaluate iron content using MRI in vivo.6 In this article, after a review of the physicochemical properties of superparamagnetic nanoparticles, the relaxation induced by these compounds will be

 2009 Jo h n Wiley & So n s, In c.

299

Advanced Review

www.wiley.com/wires/nanomed

described and the corresponding theoretical relaxation models will be presented.

SUPERPARAMAGNETIC PARTICLES Magnetite Nanoparticles Magnetite (Fe3 O4 ) belongs to the iron oxide family. It is a hard black magnetic mineral that is widespread in natural rocks. Maghemite (γ -Fe2 O3 ) is a redbrown magnetic mineral, isostructural with magnetite but with cation vacancies.7 Their global properties are quite similar, which makes it very difficult to distinguish between them. As maghemite can result from the oxidation of magnetite, a pure magnetite sample is impossible to find in our biomedical field. In the rest of this review, we will use the term maghemite, since we know that the ’magnetite’ samples surely contain a significant amount of maghemite. This does not have a tremendous influence on the relaxation properties, since the magnetic behaviors of maghemite and magnetite are practically identical. Magnetite has the particularity of containing both Fe2+ and Fe3+ ions, within an inverse spinel structure. Thirty-two oxygen anions form a facecentered cubic unit cell, with an edge length a = 0.839 nm. In this unit cell, iron ions are located on 8 tetrahedral sites (surrounded by four oxygen ions) and 16 octahedral sites (surrounded by 8 oxygen ions). The tetrahedral sites are exclusively occupied by Fe3+ ions, while Fe2+ and Fe3+ ions alternately occupy octahedral sites (Figure 1a). This organization is sometimes expressed in another formula for magnetite Fe3+ [Fe2+ Fe3+ ]O4 and maghemite Fe3+ [Fe3+ 5/3 V1/3 ]O4 where V represents a cation vacancy.

FIGURE 1 | Inverse spinel structure of

O2−

O2−

Fe3+ tetrahedral site

Fe3+ tetrahedral site

Fe2+/Fe3+ octahedral site

Fe2+/Fe3+ octahedral site

(a)

300

Magnetite is ferrimagnetic; because of superexchange oxygen-mediated coupling, all the magnetic moments of the tetrahedral iron ions are aligned in a specific direction, while all the octahedral iron magnetic moments are aligned in the opposite direction.8 Since there is the same number of octahedral and tetrahedral Fe3+ ions, they compensate for each other, and the resulting moment of a magnetite crystal arises only from the uncompensated octahedral Fe2+ ions (Figure 1b). The magnetic properties of maghemite are due to the uncompensated octahedral Fe3+ ions. The resulting magnetic moment of a magnetite (or maghemite) crystal is not directed arbitrarily; it is preferentially aligned along specific directions, called anisotropy axes. These axes are principally determined by the magneto-crystalline anisotropy field, which depends on the composition and crystallographic structure of the magnetic compound. The coupling energy between this field and the crystal magnetic moment is minimal when the magnetic moment is directed along some particular directions. For example, in the case of uniaxial anisotropy, there are only two opposite anisotropy directions and the coupling energy simply depends on the angle between the magnetic moment and this unique anisotropy axis. When the angle is equal to 0◦ or 180◦ , the energy is minimal and therefore the probability of alignment of the magnetic moment along these two directions is maximum (Figure 2). In fact, for magnetite, because of its cubic symmetry, there are eight anisotropy axes (along the diagonals of the cubic unit cell). However, for the sake of simplicity, the anisotropy of magnetite particles is often assumed to be uniaxial, that is, with a single anisotropy axis.

(b)

 2009 Jo h n Wiley & So n s, In c.

magnetite. (a) The front side of a cubic unit cell is illustrated. (b) Ferrimagnetic organization in magnetite; illustration of [1,1,1] plane.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Relaxation properties of superparamagnetic particles

0.012 µ anisotropy axis

0.010

Probability

0.008

FIGURE 2 | Model of uniaxial anisotropy for a magnetite (or

0.006

θ

0.004 0.002

maghemite) nanoparticle. The graph shows the probability of alignment of the magnetic moment in one direction with respect to the angle between this direction and the anisotropy axis (for a sphere with a radius R = 5 nm and an anisotropy constant K = 13, 500 J m−3 ).

0.000 20

0

40

60

80

100

120

140

160

180

θ (degrees)

For large ferrimagnetic crystals, not all the resulting magnetic moments are aligned. The mineral is divided into Weiss domains, within which the magnetic moments are aligned, but between which the moment’s directions are different (Figure 3a). The size of a Weiss domain in magnetite (or maghemite) is about 0.1 µm. As all maghemite contrast agents are smaller than this domain size, they must be composed of a single domain with a unique magnetic moment orientation. For these single domain crystals, a flip of the magnetic moment from one direction of anisotropy to another can be observed. If the thermal energy, given by kB T, is sufficient to overcome the anisotropy energy barrier, the magnetization will fluctuate between the

different anisotropy directions, with a characteristic time called the N´eel relaxation time.9

´ NEEL RELAXATION N´eel relaxation refers to the relaxation of the global electronic moment of a superparamagnetic crystal constituted by a ferri, ferro, or antiferromagnetic compound. It does not concern the nuclear relaxation of water protons. However, the relaxation of the particle’s huge magnetic moment will, of course, influence water proton relaxation rates. For a superparamagnetic particle of volume V, with a constant of anisotropy K, the N´eel relaxation

0.2 µm

Magnetization (Am2/kg[Fe])

100 magnetite ferritin 10

1

0.1

0.01 0.0

0.1

0.2

0.3

0.4

0.5

0.6

Typical magnetite contrast agent

Magnetic field (T)

(a)

(b)

0.7

0.8

0.9

0.10

FIGURE 3 | Illustration of Weiss domains in a large magnetite (or maghemite) crystal. (a) The small sphere represents a typical magnetite

nanoparticle, whose size is much smaller than a Weiss domain. (b) Magnetization versus magnetic field for magnetite (or maghemite) nanoparticles (plain line) with a radius R = 5 nm and a saturation magnetization Msat = 110 Am2 kg−1 (iron) and for ferritin (dashed line).

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

301

Advanced Review

www.wiley.com/wires/nanomed

time is given by:

c b

τ = τ0 e(KV)/(kb T)

where τ0 is the pre-exponential factor and kb the Boltzmann constant. The N´eel relaxation time thus increases as an exponential function of the volume of the particle. Therefore, the flipping of the magnetic moment of a magnetite crystal is observed only for nanoparticles. Indeed, for magnetite (τ0 ≈ 10−9 s, K ≈ 13500 J m−3 ), when the particle radius is about 15 nm, τ is approximately equal to 700 years, which explains why magnetite has been used for magnetic data recording. Even when the radius is larger than 10 nm, τ becomes greater than 1 ms. Because the diffusive rotation time of the particles in the colloid and the diffusion time of water molecules around the particles are far shorter than 1 ms, it seems clear that N´eel relaxation has no effect on the nuclear relaxation induced by large magnetic particles. Stable aqueous suspensions of magnetite (or maghemite) nanoparticles are said to be superparamagnetic because they show no remanence: when they are submitted to a magnetic field, their global magnetic moment aligns in the direction of the field, but when the field is set to zero, the sample magnetic moment also returns to zero. This behavior is the same as paramagnetism, but with a huge particle magnetic moment (compared to that of a paramagnetic ion, for example): this phenomenon is therefore called superparamagnetism. This nonremanence is due to the return to equilibrium of the magnetic moments through N´eel relaxation, when the particle is small enough, or through the diffusive rotation of the particle in the solvent in the case of larger particles. Macroscopically, the magnetization of a magnetite (or maghemite) nanoparticle suspension is described by a Langevin (Figure 3b) function, whose shape depends on the saturation magnetization (Msat ) and the size of the magnetite crystals. 
1 M(B) = Msat L(x) = Msat coth(x) − x µsat B with x = kT

(2)

Ferritin Ferritin is the mammal’s iron-storing protein. It is constituted by the assembly of 24 protein subunits, forming a 13-nm diameter spherical shell. Iron is stocked inside this shell in the shape of a ferrihydrite nanoparticle.10 Ferrihydrite (5 Fe2 O3 —9 H2 O) is a very common hydrated iron oxide.7 Ferritin can 302

a

(1)

Fe+3 O−2

(a)

V

V

(b)

FIGURE 4 | (a) Ferrihydrite structure. (b) Diagram of antiferro-

magnetic organization in ferrihydrite, with cation vacancies (V) and uncompensated magnetic moments (dashed arrows). (Reprinted with permission from Ref 11. Copyright 2007 AAAS).

contain up to 4500 Fe3+ ions. The number of iron atoms is called the Loading Factor (LF). Ferrihydrite has a hexagonal crystallographic structure11 and is antiferromagnetic;12 it contains two sublattices of Fe3+ ions whose magnetic moments are oriented in opposite directions (Figures 4a and 4b). The resulting magnetic moment should be null, since the two magnetic sublattices should compensate for each other. But because of surface effects (important for such nanometric compounds) and the presence of cation vacancies in the crystal, a small magnetic moment results for each ferrihydrite particle contained inside ferritin. This moment is directed along specific directions that are determined by ferrihydrite crystal anisotropy, whose constant is quite large (K = 35000 J m−3 ). This anisotropy constant and the pre-exponential factor (τ0 ≈ 10−12 s) were calculated, thanks to the measurement of the ferritin

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Relaxation properties of superparamagnetic particles

blocking temperature by different techniques,13,14 assuming a magnetic grain of 6 nm. The magnetization of ferritin is minimal compared to that of magnetite nanoparticles (Figure 3b), but given the massive presence of ferritin in some organs (liver, spleen and brain), its effect on relaxation was detected early on MR images. N´eel relaxation is also observed for ferritin (with a N´eel relaxation time of about 10−11 s), which makes it a superparamagnetic compound, with no remanence. At usual magnetic fields and at room temperature, the magnetization of a ferritin sample is simply proportional to the field.15

MAGNETIC RELAXATION INDUCED BY SUPERPARAMAGNETIC PARTICLES Maghemite Contrast Agents Main Applications in MRI Superparamagnetic magnetite particles used as contrast enhancers are classified into two main categories: a) SPIOs, small particles of iron oxide or superparamagnetic particles of iron oxide, depending on the author’s usage: particles containing several maghemite crystals within the same permeable coating. Their hydrodynamic size is often larger than 40 nm. Standard SPIOs are injected intravenously. However, some of them, intended for gastrointestinal imaging and coated within an insoluble and nonbiodegradable matrix, are administered orally. b) USPIOs, ultrasmall particles of iron oxide: particles containing a single maghemite crystal and whose hydrodynamic size is smaller than 40 nm. This type of agent is always injected intravenously.

Maghemite contrast agents (principally SPIOs) were first used 20 years ago for liver imaging: thanks to the capture of nanoparticles by Kupffer cells, an important shortening of the transverse relaxation time is observed in liver tissue.16–18 Moreover, the signal loss due to this relaxation enhancement is selective, since the maghemite particles are not internalized in lesions (void of Kupffer cells), which facilitates the early detection of tumors. The capture of the particles by the reticuloendothelial cells also allows efficient imaging of the spleen19 and of the lymph nodes.20 For applications necessitating a longer blood half-life (MR angiography, tissue perfusion imaging, Vo lu me 1, May /Ju n e 2009

functional imaging of the brain), SPIOs are not efficient, because of their elimination by the liver. USPIOs are better candidates since they stay in the blood longer and because their relaxation properties allow their use as positive contrast agents for T1 weighted imaging and angiography.21–25 Recent and promising applications of maghemite particles are related to cellular and molecular MR imaging.26–32 Bulte et al.4 defined molecular imaging as ‘the non-invasive and repetitive imaging of targeted macromolecules and biological processes’ and cellular imaging as ‘the non-invasive and repetitive imaging of targeted cells and cellular processes’. After grafting of a suited vector onto a usual contrast agent, this new type of imaging should make it possible to target specific processes such as atherosclerosis,30 apoptosis,31 and amyloid deposition in Alzheimer’s disease.32 The main problem is to bring a large quantity of contrast agent directly to the targeted zone of the body, in order to obtain a sufficient contrast during the imaging experiment. Maghemite nanoparticles are the best candidates as base elements for designing molecular or cellular imaging contrast agents: their transverse relaxivity is far greater than for any other compound and each targeted particle contains many thousands of iron atoms. The MR effect is thus optimum.

MR Relaxation Properties The efficiency of a contrast agent is often exclusively studied in aqueous solutions,33–40 even if the relaxation properties in vivo are sometimes different than those in vitro. The most important characteristics are the longitudinal and transverse relaxivities r1 and r2 of the contrast agent (see slide bar relaxivity). These quantities allow the comparison between different contrast agents and reflect their efficiency: the higher the relaxivities are, the smaller the quantity of product to be injected into the patient. Table 1 compares the relaxivities of different superparamagnetic compounds. It clearly shows that the transverse relaxivity of superparamagnetic contrast agents is far greater than their longitudinal relaxivity, which explains why they are used mainly as negative agents for T2 -weighted imaging. SPIOs are characterized by a high r2 /r1 ratio, up to 16 for AMI25 at 1.5 T. However, USPIOs present a lower r2 /r1 ratio (7.5 nm), the crystal anisotropy is great and the particle magnetic moment is locked onto the anisotropy axis. At low magnetic fields, the magnetic moment is free to flip from one anisotropy direction to the other. The time modulation of the dipolar interaction causing relaxation can have two origins: N´eel relaxation or water diffusion. This is introduced in the relaxation equations thanks to the Freed spectral density54 JF : 
 10 1 32π 2 2 NM γ µ J (ω ) = sat F 0 405 T1F
RD   3  4 1 NM 32π 2 2 γ µsat JF (ω0 ) + 2 JF (0) = 405 RD 3 T2F   1 1 + 14  2 With JF (ω) = Re 1 3 1 +  2 + 49  + 91  2 τD (4)  = iωτD + τN NM is the number of particles per volume unit, R is the magnetite crystal radius, D is the diffusion coefficient and γ the proton gyromagnetic factor. ω0 = γ B is the proton Larmor frequency. At high fields, the magnetic moment is locked onto the magnetic field direction and N´eel relaxation

u2 2

+

5u 8

u3 6

u2 8 4u4 u5 81 + 81

+

+

+

u6 648

(5)

At intermediate fields, both the mean and the fluctuating magnetic moments contribute to the relaxation induced by the superparamagnetic particles. A linear combination of equations 4 & 5 should be used. The field-dependent weighting of the two contributions reflects the respective values of the mean and fluctuating magnetic moments: 32π 2 2 1 = γ µsat T1 405

Large Particles

306

1+u+

1+

 2L(α) 2 1− − L (α) α  2 10JF (ω0 ) + L (α)9JA ( 2ω0 τD ) 
1 NM 32π 2 2 = γ µsat T2 405 RD



2 (α) [4J (ω ) + 6J (0)] 1 − 2L(α) − L F 0 F α (6) × √ +L2 (α)[4.5JA ( 2ω0 τD ) + 6JA (0)]


NM RD

 


These equations allow a perfect fitting of the field dependence curves for large particles, that is, with anisotropy energy significantly greater than thermal energy.

Small Particles For very small particles (USPIOs), low field dispersion is observed in the r1 field dependence curve (Figure 5a). Very small particles are characterized by smaller anisotropy energy so that the locking of the particle magnetization onto the anisotropy directions is attenuated. Kellar et al. developed a model discarding anisotropy energy, which results in an overestimation of the depth of the low field dip.56 Roch et al. developed a complete quantum theory introducing EA , the anisotropy energy, as a quantitative parameter of the problem.53 The predictions of that model are in good agreement with experimental observations on ultrasmall magnetite

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Relaxation properties of superparamagnetic particles

particles. However, that exact model does not provide an analytical solution for the dispersion profile, and the numerical solution of the problem requires lengthy computing time. An elegant heuristic alternative consists in calculating the relaxation rates as linear combinations of the rates calculated for EA → ∞(Eq. 6), and for EA → 056 (see further reading section). A complete description of the exact and phenomenological models can be found in Ref. 53.

Clustering and Compartmentalization The distribution of administered maghemite nanoparticles in vivo is rarely homogeneous, since they accumulate or even cluster in some specific tissues and/or cells. This affects their relaxation properties and their efficiency as contrast agents. The influence of clustering is different on longitudinal and transverse relaxation and clearly depends on the size of the clusters. The evolution of the relaxation rates during the clustering of maghemite nanoparticles was recently studied57,58 : the transverse relaxivity first increases during the agglomeration, then reaches a maximum and finally decreases, while the longitudinal relaxivity seems to decrease throughout the clustering process. This effect was already noticed in a previous study on magnetoliposomes, composed of clustered crystals with a large r2 /r1 ratio compared to SPIOs.59 The relaxation induced by maghemite nanoparticles was also shown to depend on the compartmentalization of particles within cells,60 especially in macrophages,61 monocytes,62 lymphocytes and gliosarcoma cells.63 The transverse relaxivity of SPIOs and USPIOs is two to three times lower when the magnetic particles are accumulated within cells, with a strong influence of the type of cell.62,63 This effect should be taken into account for the accurate quantification of iron-oxide-labeled cells with MRI.

Ferritin The first model proposed for ferritin-induced transverse relaxation was an adaptation of the Outer Sphere (OS) model52 already used to explain the relaxation of magnetite particles. This model assumes that the relaxation is due to the diffusion of water protons around the ferrihydrite core contained inside ferritin. The predicted transverse relaxation rate64 can be expressed as OS

r2

water diffusion coefficient, NM is the number of ferritin molecules per unit volume, C the ferritin concentration and finally µ the magnetic moment of a ferritin molecule, which is simply proportional to the magnetic field at room temperature. In fact, Eq. 7 is a combination of the classical microscopic OS theory (first term) and of the Luz and Meiboom exchange model, allowing the introduction of echo-time dependence (second term). It was previously shown64 that these two theories were equivalent in certain conditions. The predicted relaxivity of the OS model is thus dependent on the square of the magnetic field (via µ2 ) and a significant decrease in rOS 2 should be noticed for echo times smaller than τD . However, in aqueous solutions, the linearity of 1/T2 with respect to the field is well established, and the independence of the transverse relaxation on the echo time was checked on high-resolution spectrometers. These results are entirely inconsistent with the OS predictions (equation 7). Moreover, it was experimentally established65 that in aqueous solutions, the relaxation induced by ferritin—or hydrated iron oxide nanoparticles—was not due to the diffusion of water protons in the field inhomogeneities created by the ferrihydrite core (via an OS mechanism). Instead, it was shown to be caused by the exchange of protons between the core surface and bulk water (via an inner sphere mechanism). A Proton Exchange Dephasing Model (PEDM), inspired by the static dephasing regime described by Brown et al.66,67 was thus proposed for ferritin solutions.68 PEDM relaxation is caused by the irreversible dephasing of water protons because of proton exchange between bulk water and the ferrihydrite core surface. The echo time does not influence relaxation as long as it is greater than the proton exchange time, which would be a reasonable hypothesis in the case of ferrihydrite. Furthermore, the dependence of 1/T2 with the field is linear, as observed experimentally. However, this last prediction of PEDM is conditioned by the field distribution for the surface exchangeable protons, which must have a lorentzian shape. This condition has been verified for ferritin.



 τD τCP 16 µ20 γ 2 NM µ2 = 1− th 135π RD τCP τD (7)

µ0 is the vacuum magnetic permeability, τCP is the echo time, R the ferrihydrite particle radius; D the Vo lu me 1, May /Ju n e 2009

rPEDM = 2

q kB 111

(8)

q is the number of exchangeable protons, 111 is the molar concentration of protons in water and k is a constant depending on the lorentzian distribution of

 2009 Jo h n Wiley & So n s, In c.

307

Advanced Review

www.wiley.com/wires/nanomed

the magnetic field experienced by the protons on the surface of the particles. Equation 8 leads to a linear dependence of r2 with the magnetic field since the constant k is itself proportional to the magnetization of ferritin, which increases linearly with the magnetic field at room temperature. However, the increase in ferritin relaxivity observed in tissues and the effect of clustering on relaxation are not consistent with PEDM predictions. Indeed, clustering influences the relaxation regime and likely increases the OS contribution to the relaxation of ferritin in vivo. Therefore, the understanding of the relaxation induced by ferritin in vivo remains incomplete from the theoretical point of view.

CONCLUSION It is logical that superparamagnetic compounds influence water proton relaxation. However, the exact

relaxation mechanisms seem completely different for antiferromagnetic nanoparticles—like ferritin—and ferrimagnetic nanoparticles—like maghemite contrast agents. A proton exchange process dominates the relaxation of the former, while an OS process governs the relaxation of the latter. Since maghemite particles are extensively used as MRI contrast agents for conventional imaging and are promising agents for use in molecular imaging, a good understanding of their relaxation effect is needed. The influence of ferritin on in vivo MRI contrast is also crucial, since it may allow the noninvasive quantification of iron. However, understanding of the relaxation induced by superparamagnetic particles in vivo is still hazy since their distribution in tissue is not homogeneous, unlike in aqueous solutions where all these compounds have been extensively characterized. This point merits further study.

REFERENCES 1. Banci L, Bertini I, Luchinat C. Nuclear and electron relaxation. The Magnetic Nucleus-Unpaired Electron Coupling in Solution. Weinheim: Wiley-VCH; 1991.

11. Michel FM, Ehm L, Antao SM, Lee PL, Chupas PJ, et al. The structure of ferrihydrite, a nanocrystalline material. Science 2007, 316(5832):1726–1729.

2. Merbach AE, Toth E. The Chemistry of Contrast Agents in Medical Magnetic Resonance Imaging. Wiley; 2001.

12. N´eel, L. Superparamagnetisme de grains tr`es fins antiferromagnetiques. C R Acad Sci 1961 252 40754080. Translated in Selected Works of Louis N´eel (Kurti, N, ed.). New York: Gordon and Breach; 1988 107110.

3. Muller RN, Vanderelst L, Roch A, Peters JA, Csajbok E, et al. Relaxation by metal containing nanosystems. Adv Inorg Chem 2005, 57:239–292. 4. Bulte JW, Kraitchman DL. Iron oxide MR contrast agents for molecular and cellular imaging. NMR Biomed 2004, 17:484–499. 5. Gossuin Y, Muller RN, Gillis P. Relaxation induced by ferritin: a better understanding for an improved MRI iron quantification. NMR Biomed 2004, 17:427–432. 6. Brittenham GM, Badman DG. Noninvasive measurement of iron: report of an NIDDK workshop. Blood 2003, 101:15–19. 7. Cornell RM, Schwertmann U. The Iron Oxides: Structure, Properties, Reactions, Occurrences and Uses. Weinheim: Wiley-VCH; 2003. 8. N´eel L. Magnetic properties of ferrites: ferrimagnetism and antiferromagnetism. Ann Phys (Paris) 1948, 3:137–198. 9. Dormann JL. Superparamagnetism phenomenon. Rev Phys Appl 1981, 16:275–301. 10. Harrison PM, Arosio P. The ferritins: molecular properties, iron storage function and cellular regulation. Biochim Biophys Acta 1996, 1275:161–203.

308

13. Kilcoyne SH, Cywinski R. Ferritin: a model superparamagnet. J Magn Magn Mater 1995, 140–144:1466–1467. 14. Dickson DPE. Nanostructured magnetism in living systems. J Magn Magn Mater 1999, 203:46–49. 15. Gilles C, Bonville P, Wong KKW, Mann S. NonLangevin behaviour of the uncompensated magnetization in nanoparticles of artificial ferritin. Eur Phys J B 2000, 17:417–427. 16. Ros PR, Freeny PC, Harms SE, Seltzer SE, Davis PL, et al. Hepatic MR imaging with ferumoxides: a multicenter clinical trial of the safety and efficacy in the detection of focal hepatic lesions. Radiology 1995, 196:481–488. 17. Arnold P, Ward J, Wilson D, Guthrie JA, Robinson PJ. Superparamagnetic iron oxide (SPIO) enhancement in the cirrhotic liver: a comparison of two doses of ferumoxides in patients with advanced disease. Magn Reson Imaging 2003, 21:695–700. 18. Kehagias DT, Gouliamos AD, Smyrniotis V, Vlahos LJ. Diagnostic efficacy and safety of MRI of the liver with Superparamagnetic iron oxide particles (SHU555A). J Magn Reson Imaging 2001, 14:595–601.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Relaxation properties of superparamagnetic particles

19. Kreft BP, Tanimoto A, Leffler S, Finn JP, Oksendal AN, et al. Contrast-enhanced MR imaging of diffuse and focal splenic disease with use of magnetic starch microspheres. J Magn Reson Imaging 1994, 4:373–379.

32. Wadghiri YZ, Sigurdsson EM, Sadowski M, Elliott JI, Li Y, et al. Detection of Alzheimer’s amyloid in transgenic mice using magnetic resonance microimaging. Magn Reson Med 2003, 50(2):293–302.

20. Anzai Y, Blackwell KE, Hirschowitz SL, Rogers JW, Sato Y. Initial clinical experience with dextran-coated superparamagnetic iron oxide for detection of lymph node metastases in patients with head and neck cancer. Radiology 1994, 192:709–715.

33. Wang YX, Hussain SM, Krestin GP: Superparamagnetic iron oxide contrast agents: physicochemical characteristics and applications in MR imaging. Eur Radiol 2001, 11:2313–2319.

21. Bjornerud A, Johansson L. The utility of superparamagnetic contrast agents in MRI: theoretical consideration and applications in the cardiovascular system. NMR Biomed 2004, 17(7):465–477. 22. Bjerner T, Johansson L, Wikstrom G, Ericsson A, Briley-Soebo K, et al. In and ex vivo MR evaluation of acute myocardial ischemia in pigs by determining R1 in steady state after the administration of the intravascular contrast agent NC100150 injection. Invest Radiol 2004, 39(8):479–486. 23. Johansson LO, Bjerner T, Bjornerud A, Ahlstrom H, Tarlo KS, et al. Utility of NC100150 injection in cardiac MRI. Acad Radiol 2002, 9(suppl 1):S79–S81. 24. Kellar KE, Fujii DK, Gunther WH, Briley-Saebo K, Bjornerod A, et al. Important considerations in the design of iron oxide nanoparticles as contrast agents for TI-weighted MRI and MRA. Acad Radiol 2002, 9(suppl 1):S34–S37. 25. Kellar KE, Fujii DK, Gunther WH, Briley-Saebo K, Bjornerud A, et al. NC100150 Injection, a preparation of optimized iron oxide nanoparticles for positivecontrast MR angiography. J Magn Reson Imaging 2000, 11(5):488–494. 26. Bulte JW, Frank JA Imaging macrophage activity in the brain by using ultrasmall particles of iron oxide. ANJR Am J Neuroradiol 2000, 21(9):1767–1768. ¨ 27. Hartung A, Lisy MR, Herrmann K-H, Hilger I, Schuler D, et al. Labeling of macrophages using bacterial magnetosomes and their characterization by magnetic resonance imaging. J Magn Magn Mater 2007, 311:454–459. 28. Becker C, Hodenius M, Blendinger G, Sechi A, Hieronymus T, et al. Uptake of magnetic nanoparticles into cells for cell tracking. J Magn Magn Mater 2007, 311:234–237.

34. Bulte JW, Vymazal J, Brooks RA, Pierpaoli C, Frank JA. Frequency dependence of MR relaxation times. II. Iron oxides. J Magn Reson Imaging 1993, 3(4):641–648. 35. Jung CW, Jacobs P. Physical and chemical properties of Superparamagnetic iron oxide MR contrast agents: ferumoxides, ferumoxtran, ferumoxsil. Magn Reson Imaging 1995, 13:661–674. 36. Bulte JW, Brooks RA, Moskowitz BM, Bryant LH, Frank JA. Relaxometry and magnetometry of the MR contrast agent MION-46 L. Magn Reson Med 1999, 42:379–384. 37. Bulte JW, Brooks RA, Moskowitz BM, Bryant LH Jr, Frank JA. T1 and T2 relaxometry of monocrystalline iron oxide nanoparticles (MION-46L): theory and experiment. Acad Radiol 1998, 5(suppl 1):S137–S140, discussion S145–S146. 38. Koenig SH, Kellar KE, Fujii DK, Gunther WH, BrileySaebo K, et al. Three types of physical measurements needed to characterize iron oxide nanoparticles for MRI and MRA: magnetization, relaxometry, and light scattering. Acad Radiol 2002, 9(suppl 1):S5–S10. 39. Bjornerud A, Wendland M, Johansson L, Ahlstrom H, Higgins C, et al. Use of intravascular contrast agents in MRI. Acad Radiol 1998, 5(suppl 1):223–225. 40. Bulte JW, Douglas T, Mann S, Frankel RB, Moskowitz BM, et al. Magnetoferritin: characterization of a novel superparamagnetic MR contrast agent. J Magn Reson Imaging 1994, 4(3):497–505. 41. Gossuin Y, Roch A, Muller RN, Gillis P. Relaxation induced by ferritin and ferritin-like magnetic particles: the role of proton exchange. Magn Reson Med 2000, 43:237–243. 42. Vymazal J, Brooks RA, Zak O, McRill C, Shen C, et al. T1 and T2 of ferritin at different field strengths: effect on MRI. Magn Reson Med 1992, 27:368–374.

29. Bulte JW, Douglas T, Witwer B, Zhang S-C, Strable E, et al. Magnetodendrimers allow endosomal magnetic labeling and in vivo tracking of stem cells. Nat Biotechnol 2001, 19:1141–1147.

43. Vymazal J, Brooks RA, Baumgarner C, Tran V, Katz D, et al. The relation between brain iron and NMR relaxation times: an in vitro study. Magn Reson Med 1996, 35(1):56–61.

30. Jaffer F, Nahrendorf M, Sosnovik D, Kelly KA, Aikawa E, et al. Cellular imaging of inflammation in atherosclerosis using magnetofluorescent nanomaterials. Mol Imaging 2006, 2(5):85–92.

44. Papakonstantinou OG, Maris TG, Kostaridou V, Gouliamos AD, Koutoulas GK, et al. Assessment of liver iron overload by T2-quantitative magnetic resonance imaging: correlation of T2-QMRI measurements with serum ferritin concentration and histologic grading of siderosis. Magn Reson Imaging 1995, 13:967–977.

31. Zhao M, Beauregard DA, Loizou L, Davletov B, Brindle KM. Non-invasive detection of apoptosis using magnetic resonance imaging and a targeted contrast agent. Nat Med 2001, 7:1241–1244.

Vo lu me 1, May /Ju n e 2009

45. Ma HL, Xu YF, Qi XR, Maitani Y, Nagai T. Superparamagnetic iron oxide nanoparticles stabilized by

 2009 Jo h n Wiley & So n s, In c.

309

Advanced Review

www.wiley.com/wires/nanomed

alginate: pharmacokinetics, tissue distribution, and applications in detecting liver cancers. Int J Pharm 2008, 354(1–2):217–226. 46. Lee JH, Huh YM, Jun Y, Seo J, Jang J, et al. Artificially engineered magnetic nanoparticles for ultra-sensitive molecular imaging. Nat Med 2007, 13(1):95–99. 47. Wan J, Cai W, Meng X, Liu E. Monodisperse watersoluble magnetite nanoparticles prepared by polyol process for high-performance magnetic resonance imaging. Chem Commun 2007, 47:5004–5006. 48. Gossuin Y, Roch A, Lo Bue F, Muller RN, Gillis P. Nuclear magnetic relaxation dispersion of ferritin and ferritin-like magnetic particle solutions: a pH-effect study. Magn Reson Med 2001, 46(3):476–481. 49. Vymazal J, Zak O, Bulte JWM, Aisen P, Brooks RA. T1 and T2 of ferritin solutions: effect of loading factor. Magn Reson Med 1996, 36(1):61–65. 50. Gossuin Y, Burtea C, Monseux A, Toubeau G, Roch A, et al. Ferritin-induced relaxation in tissues: an in vitro study. J Magn Reson Imaging 2004, 20:690–696. 51. Gossuin Y, Gillis P, Muller RN, Hocq A. Relaxation by clustered ferritin: a model for ferritin-induced relaxation in vivo. NMR Biomed 2007, 20(8):749–756. 52. Gillis P, Koenig SH. Transverse relaxation of solvent protons induced by magnetized spheres: application to ferritin, erythrocytes, and magnetite. Magn Reson Med 1987, 5(4):323–345. 53. Roch A, Muller RN, Gillis P. Theory of proton relaxation induced by superparamagnetic particles. J Chem Phys 1999, 110:5403–5411. 54. Freed JH. Dynamic effects of pair correlation functions on spin relaxation by translational diffusion in liquids. II. Finite jumps and independent T1 processes. J Chem Phys 1978, 9:4034–4037. 55. Ayant Y, Belorizky E, Alizon J, Gallice J. Calculation of spectral density resulting from random translational movement with relaxation by magnetic dipolar interaction in liquids. J Phys A 1975, 36:991–1004.

58. Larsen BA, Haag MA, Serkova NJ, Shroyer KR, Stoldt CR. Controlled aggregation of superparamagnetic iron oxide nanoparticles for the development of molecular magnetic resonance imaging probes. Nanotechnology 2008, 19:265102. 59. Bulte JW, de Cuyper M, Despres D, Frank JA. Preparation, relaxometry, and biokinetics of PEGylated magnetoliposomes as MR contrast agent. J Magn Magn Mater 1999, 194:204–209. 60. Bowen CV, Zhang X, Saab G, Gareau PJ, Rutt BK. Application of the static Dephasing regime theory to superparamagnetic iron-oxide loaded cells. Magn Reson Med 2002, 48:52–61. 61. Hartrung A, Lisy MR, Herrmann K-H, Hilger I, ¨ Schuler D, et al. Labeling of macrophages using bacterial magnetosomes and their characterization by magnetic resonance imaging. J Magn Magn Mater 2007, 311:454–459. 62. Rad AM, Arbab AS, Iskander ASM, Jiang Q, SoltanianZadeh H. Quantification of superparamagnetic iron oxide (SPIO)-labeled cells using MRI. J Magn Reson Imaging 2007, 26:366–374. 63. Simon GH, Bauer J, Saborovski O, Fu Y, Corot C, et al. T1 and T2 relaxivity of intracellular and extracellular USPIO at 1.5T and 3T clinical MR scanning. Eur Radiol 2006, 16:738–745. 64. Brooks RA, Moiny F, Gillis P. On T2-shortening by weakly magnetized particles: the chemical exchange model. Magn Reson Med 2001, 45:1014–1020. 65. Gossuin Y, Roch A, Muller RN, Gillis P. An evaluation of the contributions of diffusion and exchange in relaxation enhancement by MRI contrast agents. J Magn Reson 2002, 158(1–2):36–42. 66. Brown RJS. Distribution of fields from randomly placed dipoles: free-precession signal decay as result of magnetite grains. Phys Rev 1961, 121:1379–1382.

56. Koenig SH, Kellar KE. Theory of 1/T1 and 1/T2 NMRD profiles of solutions of magnetic nanoparticles. Magn Reson Med 1995, 34(2):227–233.

67. Yablonskiy DA, Haacke EM. Theory of NMR signal behavior in magnetically inhomogeneous tissues: the static dephasing regime. Magn Reson Med 1994, 32:749–763.

57. Roch A, Gossuin Y, Muller RN, Gillis P. Superparamagnetic colloid suspensions: water magnetic relaxation and clustering. J Magn Magn Mater 2005, 293:532–539.

68. Gossuin Y, Roch A, Muller RN, Gillis P, LoBlue F. Anomalous nuclear magnetic relaxation of aqueous solutions of ferritin: an unprecedented first-order mechanism. Magn Reson Med 2002, 48(6):959–964.

FURTHER READING http://www.mr-tip.com/serv1.php?type=coa for interesting data on contrast agents. http://www.cerm.unifi.it/relax/relaxometry2.html for the fitting program.

310

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

Advanced Review

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis Shelton D. Caruthers,1∗ Tillmann Cyrus,2 Patrick M. Winter,2 Samuel A. Wickline2 and Gregory M. Lanza2 Complementary developments in nanotechnology, genomics, proteomics, molecular biology and imaging offer the potential for early, accurate diagnosis. Molecularly-targeted diagnostic imaging agents will allow noninvasive phenotypic characterization of pathologies and, therefore, tailored treatment close to the onset. For atherosclerosis, this includes anti-angiogenic therapy with specificallytargeted drug delivery systems to arrest the development of plaques before they impinge upon the lumen. Additionally, monitoring the application and effects of this targeted therapy in a serial fashion will be important. This review covers the specific application of αν β3 -targeted anti-angiogenic perfluorocarbon nanoparticles in (1) the detection of molecular markers for atherosclerosis, (2) the immediate verification of drug delivery with image-based prediction of therapy outcomes, and (3) the serial, noninvasive observation of therapeutic efficacy.  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 311–323

A

therosclerosis, a slow and complex process, is clearly the major pathological component of cardiovascular disease, lurking quietly in a large majority of the population.1–5 With little warning, quiescent disease can destabilize with sudden and catastrophic results.6 As with all medicine, the goals of improving disease management and quality of life hinge on improving not only therapy, but also diagnosis. For atherosclerosis, one key to preventing myocardial infarction or stroke is to detect and quantify, specifically and sensitively, those plaques which are more apt to rupture versus those which are, and will remain, stable. Moreover, the early, specific detection of the atherosclerotic process will allow different, and perhaps more however, moderate, therapy to be more effective in reducing the overall atherosclerotic burden. Many nanomedical applications, as evidenced in this publication, are emerging to address each of these areas;7–12 the majority remained focused on detection and have not yet evolved to treatment. This review is not intended to be a comprehensive review of

∗ Correspondence

to: [email protected]

1

Washington University School of Medicine and Philips Medical Systems, St. Louis, MO, USA 2 Washington

University School of Medicine, St. Louis, MO, USA

DOI: 10.1002/wnan.009

Vo lu me 1, May /Ju n e 2009

nanomedical applications in atherosclerosis,8,11,13–18 but instead it presents recent experiences with a very specific nanoparticle application—site-targeted, liquid perfluorcarbon emulsions—in the diagnosis, therapy, and serial monitoring of therapy in atherosclerosis.

DIAGNOSING ATHEROSCLEROSIS WITH TARGETED NANOPARTICLES Pathology and Biomarkers of Atherosclerosis Significant atherosclerotic disease may exist without compromising the arterial lumen, and many acute coronary syndromes are the result of sudden lumenal thrombosis, which forms on a ruptured or eroded plaque.19–22 The disruption of the endothelial layer because of plaque erosion or focal rupture causes fibrin deposition making this process amenable to molecularly targeted imaging techniques.20,23 Furthermore, the plaque prone to erosion may have endothelial cell damage resulting in increased platelet adhesion over a period of time prior to complete erosion, providing a window of opportunity for detection before vascular complications ensue. In contrast to conventional angiography, which determines lumenal obstruction by contrast deficits,

 2009 Jo h n Wiley & So n s, In c.

311

Advanced Review

www.wiley.com/wires/nanomed

molecularly targeted nanoparticulate agents can interrogate the vessel from the lumenal side, detecting fibrin, or they can approach through vasa vasorum in the outer vascular wall to reach targets, such as the αν β3 integrin,24,25 in angiogenic vessels that are feeding the developing plaque and promoting intraplaque hemorrhage.26 Other nanoparticles are disguised as high density lipoproteins (HDL)11 or target to macrophages,27,28 both key components in the atherosclerotic process. Thus, molecularly targeted nanoparticles offer numerous opportunities to image atherosclerotic plaque and noninvasively provide risk assessments regarding the vulnerability of individual atheroma. Fibrin represents a rich biomarker target for targeted agents to facilitate sensitive, specific detection and, potentially, quantification of exposed microthrombi resulting from atherosclerotic plaque erosion or rupture.29 Despite its value as a specific marker for plaque vulnerable to further rupture, fibrin is exposed to the vessel lumen relatively late in the atherosclerotic process—in many cases only at the first, but catastrophic, rupture. For this reason, additional markers of early atherosclerosis are valuable. While an ever-increasing number of such markers exist, the αν β3 integrin is interesting because of its association with angiogenesis. At the distal growth zone of angiogenic vasculature, the αν β3 integrin is transiently expressed during the growth phase allowing for specific targeting of this integrin which is otherwise not expressed on physiologically intact endothelium.30,31 Importantly, antagonists to the αν β3 integrin promote tumor regression in animal32 and human30 tumor models via inhibition of angiogenesis. The integrins such as αν β3 and α5 β1 , another well characterized target for molecular imaging,33 may provide viable imaging targets. Vascular endothelial growth factor (VEGF) is an important angiogenic factor induced by local hypoxia and interacts with tyrosine kinase receptors which may also be available for targeted imaging.34 As demonstrated in the chapters of this Interdisciplinary Reviews, these and other biomarkers are the object of many nanotechnology-based molecularly targeted imaging agents.

Nanoparticulate Emulsions of Liquid Perfluorocarbon Nano-sized agents are uniquely sized to offer improved efficacy of molecularly targeted contrast agents.13,35,36 The particles can be made large enough to be constrained within the intact vasculature, so that unlike typical molecular contrast agents they 312

do not extravasate perfusing the entire body; yet they are small enough to reach targets of interest. Furthermore, the increased surface-area-to-volume ratio provides ample space to affix multiple copies of the targeting moiety and/or the contrast-creating component [e.g., thousands of Gd chelates per nanoparticle for magnetic resonance imaging (MRI) contrast]. The nanometer scale of these agents allows appropriate homing with high target affinity while highly concentrating the imaging component in order for sparse biomarkers to be imaged with relatively insensitive imaging techniques. This review focuses on liquid perfluorocarbon-based nanoparticle agents targeted to the angiogenic processes of atherosclerosis. The formulation of this particular contrast agent is described elsewhere37 (including other sections of this overall publication) and is beyond the scope of this specific review. In brief, the emulsion comprises liquid perfluorocarbon encapsulated within a phospholipid monolayer and dispersed in water. Different perfluorocarbons can be used for the core, including perfluorodichlorooctane, perfluorodecaline, perfluoro 15-crown-5 ether (CE), and, most commonly, perfluorooctyl bromide (PFOB). The use of perfluorocarbons has been explored for other medical applications, including partial liquid lung ventilation,38–40 gastrointestinal X-ray contrast agent,41,42 and blood substitutes.43–48 The phospholipid-wrapped core of these emulsions varies in size from 200 to 300 nm in diameter, depending on, among other things, the perfluorocarbon used. For PFOB, the particles are typically 250 nm making them much smaller than microbubbles, but large enough that they are restricted within the intact vasculature. Furthermore, the surface area to volume ratio is extremely high providing ample space to attach, for example, homing ligands. Additionally, imaging agents, such as chelates of paramagnetic or radioactive elements, can be attached to further functionalize the particles. This generalized paradigm of a ‘skeleton’ onto which targeting, imaging, and even therapeutic agents can be attached (Figure 1) is employed in many different nanomedical constructs.37,49–56

Noninvasive Imaging of Site-Targeted Contrast Agents One of the earliest demonstrations of imaging molecularly targeted liquid perfluorocarbon emulsions was in 1996.37 In this study, an emulsion was targeted to thrombi via biotin and avidin interactions58–60 using an antifibrin monoclonal antibody.61,62 Demonstrated both in vitro with fibrin-rich clots formed on

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

FIGURE 1 | Generalized paradigm for targeted nanoparticle

contrast agents. The nanoparticle core provides a framework onto which volumes of the targeting system, imaging and therapeutic agents can be placed. Nanoscaled agents, with their high surface-area-to-volume ratio, provide the amplification strategies that permit molecular imaging to be feasible when targeting sparse epitopes. (Reprinted, with permission, from Ref. 57. Copyright 2003.)

a suture and in vivo with femoral artery thrombus formed by a transmural electrode, this site-targeted agent was imaged using ultrasound (7.5 MHz) causing marked increase in acoustic reflectivity versus the control thrombi which lacked contrast agent (Figure 2). Unlike that for gas-filled microbubbles, the sensitivity of ultrasound for these smaller, liquid nanoparticles is relatively low, thus rendering the agent difficult to detect by ultrasound until it is bound and accumulated at the site of interest. The echogenic contrast mechanisms have been well described,63–68 and novel reconstruction algorithms69,70 have been developed to further improve both detectability and quantification of the bound agent. For use with magnetic resonance imaging (MRI) the acoustic agent was further modified such that the outer phospholipid monolayer incorporated multiple gadolinium chelates.71 In contrast to previous attempts of attaching only a few gadolinium ions to a single targeting ligand (a technique that works well with radionuclides), the nanotechnology approach—which, exploiting the vast surface area, includes thousands of copies per targeted particle—provides a mechanism to amplify greatly the imaging effect,72 thus making it possible for minute concentrations (i.e., picomolar) of targeted agent to be visualized conspicuously via standard T1 -weighted MRI.73 Akin to the previous ultrasound experiments, this ultraparamagnetic MRI agent was used to target and image fibrin-rich clots in vitro,74 thrombi in canine jugular veins (Figure 3), and exposed

fibrin within microfissures in human carotid artery specimens taken from symptomatic patients.75 As discussed above, fibrin represents a rich biomarker target for these agents to facilitate sensitive, specific detection and, potentially, quantification of exposed microthrombi resulting from atherosclerotic plaque erosion or rupture.29 Despite its value as a specific marker for plaque vulnerable to further rupture, fibrin is exposed to the vessel lumen relatively late in the atherosclerotic process—in many cases too late for treatment. For this reason, other markers, present in earlier stages of atherosclerosis, are valuable. The αν β3 -integrin is interesting because, as a result of its association with angiogenesis, it provides a target not only for detection of early atherosclerosis but also for delivery of anti-angiogenic therapy. Employing paramagnetic nanoparticles functionalized to target neovasculature through the αν β3 integin,76 Winter et al.,77 used MRI and a hyperlipidemic rabbit model to visualize early atherosclerosis. (When fed 0.5–1.0% cholesterol diets, these New Zealand white rabbits develop early neovascular expansion of the vasa vasorum which fuels plaque progression.78,79 ) To cholesterol-fed rabbits and to healthy control animals, αν β3 -targeted paramagnetic nanoparticle emulsions were administered intravenously. Additionally, nontargeted paramagnetic nanoparticle emulsion was given to a cholesterolfed group. Imaging was performed from just prior to injection through 2 h postinjection. From the highresolution, T1 -weighted, black-blood MR images, the Before contrast

FIGURE 2 | Ultrasound imaging of femoral artery thrombus

before (a) and after (b) exposure to targeted perfluorocarbon nanoparticles. (a) While the transmural electrode (anode) and the vessel wall boundaries of the femoral artery are clearly delineated with a 7.5 MHz linear-array, focused transducer, the acute thrombus is poorly visualized. (b) After exposure to the targeted emulsion, the thrombus is easily visualized. (Reprinted, with permission, from Ref. 37. Copyright 1996.)

Vo lu me 1, May /Ju n e 2009

(a)

After contrast

(b)

Anode

 2009 Jo h n Wiley & So n s, In c.

Thrombus

313

Advanced Review

www.wiley.com/wires/nanomed

(a)

(b)

(c)

Jugular vein

0.5 µm

0.5 µm

FIGURE 3 | Fibrin-targeted paramagnetic nanoparticles. A scanning electron micrograph (a) shows the fibrin tendrils within a clot. Targeted

nanoparticles densely decorate the fibrin (b), bringing large volumes of imaging agent (gadolinium chelates) per binding site. The effect, in this canine model (c), is a marked enhancement (arrow) as compared to the control clot in the contralateral vein. (Reprinted, with permission, from Ref. 75. Copyright 2001.)

entire aortic wall, from diaphragm to renal arteries, was segmented80 and analyzed for signal enhancement. Regional analysis showed patchy areas of generous signal enhancement because of the accumulated binding of the αν β3 -targeted paramagnetic nanoparticles (Figure 4). Likewise, considering the entire aorta, the atherosclerotic group given the αν β3 targeted paramagnetic agent exhibited significant signal enhancement as compared to the two control groups. Histology confirmed the colocalization of neovasculature associated with atherosclerotic plaques and the dearth of microvessels in healthy aortas. To test the specificity of the targeted agent, a competitive blockade experiment was performed. Prior to imaging, atherosclerotic rabbits received a dose of high-avidity, αν β3 -targeted nanoparticles that lacked gadolinium (i.e., MR invisible) to occupy binding sites. Following this, αν β3 -targeted paramagnetic nanoparticles were given and imaged as before but resulted in greatly reduced signal enhancement as compared to the noncompetition αν β3 -targeted group. This confirms that the T1 -weighted MR signal enhancement which was seen in the hyperlipidemic rabbits was due

to the specific binding of the paramagnetic agent to the αν β3 integrin associated with angiogenesis arising from early-stage atherosclerosis.

Quantitative Imaging of Multiple Biomarkers To asses the severity of atherosclerosis, the quantification of biomarkers such as microvascular angiogenesis and exposed fibrin could be useful.6,81–84 Evaluating ultrasound reflectivity can quantify sparsely bound perfluorocarbon agents.67 Likewise, comparing signal intensity changes because of gadolinium on T1 -weighted MRI is another relative measure of such quantification. However, more precise measures can be derived from more traditional quantitative imaging such as T1 mapping,85 etc. MRI contrast agents interact with their immediate milieu effecting changes in both T1 and T2 relaxation phenomena of nearby hydrogen nuclei. For the ideal situation and known contrast agent relaxivities, these effects can be modeled via the Bloch equation86,87 to predict signal intensity as a function of agent concentration for a given pulse sequence. Employing these

Slice 1

10

20

30 Diaphragm

Renal

100 80 60 40 Pre

314

Post

Segmented

Enhancement

20

FIGURE 4 | In vivo MR images of hyperlipidemic rabbit aorta. (a) The αν β3 -targeted nanoparticles cause heterogeneous signal enhancement along the entire vessel wall. (b) Transverse images before (pre) and after (post) nanoparticles, after segmentation of the vessel wall (segmented), and the final images (enhancement) with color encoding of the percent enhancement resulting from the contrast agent bound to the marker of angiogenesis in this model of early atherosclerosis. (Reprinted, with permission, from Ref. 77. Copyright 2003.)

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

methods, Morawski et al.73 not only modeled the minimum concentrations of paramagnetic perfluorocarbon nanoparticles required to be bound within an imaging voxel to produce conspicuous signal enhancement, but also calculated, in an in vitro experiment, the number of targeted nanoparticles bound to cells in culture. This algorithm, while more precise than simple T1 -weighted signal comparisons, is still limited by assumptions about relaxivities, water exchange rates, and other confounding effects. As an alternative to modeling or measuring indirect effects of contrast agents, one could, as is the case for nuclear medicine radioisotopes, measure the signal of the contrast agent directly. While not all nuclei exhibit the magnetic resonance effect and therefore cannot be measured directly, fluorine-19, like hydrogen-1, has one unpaired proton and no unpaired neutrons thus with a net spin of 1/2 exhibits the nuclear magnetic resonance (NMR) phenomena. Serendipitously, the gyromagnetic ratio (γ ) for 19 F is also close to that of 1 H (i.e., 40.1 MHz/T versus 42.6 MHz/T, respectively) which means not only are their resonance frequencies relatively similar but so are their relative signal strengths, with 19 F being approximately 83% that of 1 H. Furthermore, while the 19 F isotope of fluorine has a natural abundance of near 100%, the biological presence is virtually zero. Therefore, the 19 F nuclei which are highly concentrated within the perfluorocarbon core of the nanoparticle emulsions are a prime candidate for direct MR spectroscopy and imaging without surrounding signal from endogenous fluorine. Unlike hydrogen with only one electron, fluorine, with nine, is sensitive to its environment through chemical shift to a much greater extent. This is the key reason such a broad variety of applications for 19 F MR

spectroscopy and imaging has arisen, including the mapping of physiological pO2 tension,88–90 the study of tumor metabolism,91–93 and the characterization of liquid ventilation with perfluorocarbons.94,95 Using the 19 F signal at 4.7T from fibrintargeted, paramagnetic perfluorocarbon nanoparticles, Morawski et al.96 accurately quantified the varying amounts of contrast agents bound to in vitro clots, verifying the results with neutron activation analysis97 of the gadolinium on the nanoparticles. Furthermore, this technique was applied to human endarterectomy samples, providing a quantitative measure of the amount of exposed fibrin associated with atherosclerotic plaques. Extending this to a 1.5T MR system using rapid steady-state imaging techniques,98 in vitro fibrin-bound nanoparticles of two different species of perfluorocarbon (PFOB and CE) in varying volumes were not only quantified via clinically relevant imaging techniques but were also identified and independently imaged on the basis of their unique MR spectral signatures (see Figure 5). From this demonstration on fibrin clots, one might extrapolate the potential of using multiple perfluorocarbon nanoparticle agents, each targeted to a different epitope, to perform noninvasive, imaging-based immunohistochemistry, e.g., quantifying simultaneously the amount of angiogenesis and exposed fibrin associated with an atherosclerotic plaque as an indicator of its pathophysiological significance. Of course, the whole raison d’ˆetre of more accurate, quantifiable, early diagnosis is that appropriate therapy can be applied.99 To some extent, having an early diagnosis of silent lesions is enough to allow current therapy to be applied more effectively. However, the more optimistic approach includes novel (a)

19

F Spectra

(b)

19F

Images

FIGURE 5 | 19 F MR spectroscopy and imaging of liquid

perfluorocarbon nanoparticles bound to in vitro fibrin clots. The four spectra (a) acquired from the corresponding four clots (b) demonstrate the changing concentrations of the two different nanoparticle species applied. The three columns of 19 F images (b), which have no proton background, represent three ‘weightings’ of the same four clots. Without selectivity (NS), all clots have high signal. Employing a species-selective excitation allows independent visualization of the bound 15-crown-5 ether (CE) or perfluorooctyl bromide (PFOB) nanoparticles. (Reprinted, with permission, from Ref. 36. Copyright 2006 Springer.) 19 F

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

CE PF OB NS

CE

PF OB

315

Advanced Review

www.wiley.com/wires/nanomed

methods of therapy designed in collusion with the targeted diagnostics to attack subclinical diseases presymptomatically.100–102

DELIVERING ANTI-ANGIOGENIC THERAPY WITH TARGETED NANOPARTICLES Angiogenesis is Required for Atherosclerosis As discussed above, angiogenesis of the vasa vasorum is required for the progression of atherosclerosis. While this neovasculature provides a flourishing biomarker for detecting early atherosclerosis, it also acts as a pivotal point upon which therapy can be applied to arrest atherosclerosis.103 Jain et al.26 propose that, in a process similar to anti-angiogenic tumor therapy, the ‘pruning and normalization’ of immature intraplaque vessels help prevent intraplaque hemorrhage thereby reducing the vulnerability of the plaque to rupture. Moulton et al.104 showed that anti-angiogenic therapy reduces the rate of plaque growth in the apolipoprotein E insufficient mouse model of atherosclerosis. Administering the antiangiogenic drug TNP-470,105 a water soluble analog of fumagillin, at 30 mg/kg every other day for 16 weeks inhibited the plaque growth by 70%.

Anti-Angiogenic Drug Delivery with Targeted Nanoparticles Along with the ability to target and image specific molecular biomarkers with perfluorocarbon nanoparticles comes the unique opportunity also to deliver and concentrate potent individualized therapy directly to disease sites.8,14,100,106–108 By dissolving lipophilic drugs into the phospholipid monolayer of the nanoparticle core or by affixing lipophobic drugs onto the outer surface with a lipophilic anchor, the diagnostic agent is transformed into a therapeutic delivery device100

while maintaining its properties as an imaging agent. This approach is not an ‘extended release’ mechanism, nor does it require particle destruction or cellular internalization to deliver the drug. Rather, the drug is retained with the nanoparticle core and, only upon ligand-directed binding, is conveyed to the cells by ‘contact facilitated drug delivery,’ wherein lipids and the associated drugs are transferred between the nanoparticle monolayer and the proximate cell membrane109 (Figure 6). Normally, for a circulating lipid particle system, collision-mediated lipid–lipid exchange would be inconsequential110–112 ; but for a tethered system like the targeted nanoparticles, the exchange becomes significant. This mechanism of targeted drug delivery with the perfluorocarbon nanoparticles was first demonstrated with particles targeted to smooth muscle cells (via tissue factor) delivering antiproliferative therapy (doxorubicin or paclitaxel).100 These studies showed that through ‘contact-facilitated drug delivery’ the drugs had an antiproliferative effect, regardless of dose (0.2 or 2.0 mol %). However, when untargeted, the drug-laden nanoparticles had (for doxorubicin) significantly reduced or (for paclitaxel) no effect on cell proliferation. On the basis of the diagnostic perfluorocarbon nanoparticle emulsions which were targeted to angiogenesis through anti-αν β3 peptidomimetics, a therapeutic anti-angiogenic nanoparticle emulsion was created by replacing some of the lecithin in the outer core surfactant with fumagillin (0.2 mol%).113,114 Owing to its lipophilicity, the fumagillin was well retained in the monolayer. Under sink conditions of in vitro dialysis, less than 9% of the total incorporated drug was released with the majority

Contact facilitated drug delivery (a)

(b)

Circulate to target

316

(c)

Bind to biomarker

Lipid-drug exchange

FIGURE 6 | Schematic representation of ‘contact facilitated drug delivery.’ Phospholipids and drug within the perfluorocarbon nanoparticle surfactant exchange with lipids of the target membrane through a convection process rather than diffusion, as is common among other targeted systems. Without the stable membrane contact, enabled through specific binding, drug is not released from the particle. (Reprinted, with permission, from Ref. 106. Copyright 2004.)

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

(ca. 5%) in the first 24 h and no further detectable dissociation after the third day. Returning to the cholesterol-fed rabbit model of early atherosclerosis, Winter et al.114 dosed groups with one of three emulsions: αν β3 -targeted fumagillin nanoparticles, αν β3 -targeted nanoparticles without fumagillin, and nontargeted fumagillin nanoparticles. All three nanoparticle types included gadolinium so that MRI could be performed at treatment not only to asses the level of vasa vasorum neovasculature as indicated by αν β3 integrin expression but also to confirm and quantify the specific delivery of drug-laden nanoparticles, if targeted. One week later, all subjects were reimaged using the diagnostic αν β3 -targeted paramagnetic nanoparticles (i.e., no drug). The group which received the targeted anti-angiogenic nanoparticles exhibited a significant reduction in both the spatial distribution and level of αν β3 -related signal enhancement, whereas the groups which received no drug or nontargeted drug-carrying nanoparticles exhibited no therapeutic effect (Figure 7). These image-based findings of anti-angiogenic therapy were corroborated αnβ3-Targeted without Drug

αnβ3-Targeted with Drug

50 Treatment

10

1 Wk Post

FIGURE 7 | Molecular imaging of therapy response to

anti-angiogenic nanoparticles. MR imaging of hyperlipidemic rabbit aorta (as in Figure 4) showing the false-colored overlay of signal enhancement as a result of αν β3 targeted paramagnetic nanoparticles at time of treatment (top row). On follow-up imaging 1 week later (bottom row), the subject receiving drug-laden nanoparticles (left) shows marked reduction in the atherosclerosis-related angiogenesis compared to the subject receiving no drug (right). (Reprinted, with permission, from Ref. 114. Copyright 2006.)

Vo lu me 1, May /Ju n e 2009

by histology. Histology confirmed heterogeneously distributed intimal thickening consistent with early atherosclerosis in all groups, but the amount of neovasculature, which was concentrated to areas of the adventitia juxtaposed to the thickened intima, was significantly less in those given αν β3 -targeted anti-angiogenic nanoparticles than in those given drug-free particles. In these experiments, an important finding is that the T1 -weighted MR image signal enhancement in the wall of the aorta on the day of initial treatment with the αν β3 -targeted anti-angiogenic paramagnetic nanoparticles correlated with the extent of therapeutic response on 1-week follow-up imaging. In other words, at the time of dosing, noninvasive imaging not only quantitatively confirmed the targeted delivery of drug but also predicted the effect. This delivery confirmation is an important aspect of image-guided monitoring of targeted therapeutics.

IMAGE-GUIDED MONITORING OF THERAPY As site-targeted nanotherapeutics have little effect unless binding occurs, the validation, localization, and quantification of drug delivery are important. With multifunctional agents such as the anti-angiogenic paramagnetic nanoparticle emulsion, this monitoring of therapy can be performed noninvasively through tomographic imaging. Hence clinically-relevant quantitative imaging, as discussed above, will be important for initial dose measurements as well as serial comparisons of response. In the above example,114 imaging upon treatment was used to verify the delivery of drug-laden nanoparticles to the sites of angiogenesis associated with neointimal thickening. But beyond simple confirmation, they showed that the areas exhibiting greatest signal enhancement at the time of dosing (i.e. receiving the most nanoparticles) were the areas that changed the most after treatment (unlike in the control group receiving drug-free nanoparticles which remained unchanged). In addition to imaging the initial dose of antiangiogenic nanoparticles, imaging with the drugfree (or drug-laden) nanoparticles offers a tool for rapid, noninvasive, serial follow-up monitoring of treatment response. Though initially this started with a single, 1-week follow-up of therapy, this approach has since been extended to longer-term serial monitoring of therapy.115 Again in the model of early atherosclerosis, Winter et al.116 used the diagnostic αν β3 -targeted nanoparticles weekly to monitor the

 2009 Jo h n Wiley & So n s, In c.

317

Advanced Review

www.wiley.com/wires/nanomed

effect of the targeted anti-angiogenic nanoparticles demonstrating that after giving a single dose, discontinuing therapy but continuing the high-cholesterol diet, the disease-related angiogenesis recrudesced over a 4 week period. Extending this even longer, this group117 used nanoparticle-based, image-guided monitoring to follow for 8 weeks the effects of combination therapy, wherein oral statins were combined with αν β3 -targeted anti-angiogenic nanoparticles. With a temporal resolution of 1 week, they were able to follow the increases and decreases of the biomarker for angiogenesis. While this temporal precision is required for academic and preclinical studies, the clinical scenario for image-guided monitoring of therapy may be a bit different while still affording a significant improvement over the months-long waits patients currently must endure for evidence of a response (or lack thereof)—a point of particular import for, e.g., chemotherapy.

CONCLUSION Synergistically combining targeted drug delivery and molecular imaging, angiogenesis-targeted perfluorocarbon nanoparticle emulsions offer the potential to revolutionize the detection and treatment of cardiovascular disease. Homed to specific molecular epitopes and extended to a multispectral palette, they promise noninvasive, phenotypic characterization of pathology with implications toward tailored therapy. Drug delivery agents that, with noninvasive imaging techniques, are detectable and quantifiable at the targeted site may ultimately give instant verification of delivery plus confirmation of therapeutic efficacy through serial characterization of the molecular epitope expression.

REFERENCES 1. Pearson TA. Coronary arteriography in the study of the epidemiology of coronary artery disease. Epidemiol Rev 1984, 6:140–166. 2. Kiechl S, Willeit J. The natural course of atherosclerosis. Part I: incidence and progression. Arterioscler Thromb Vasc Biol 1999, 19(6):1484–1490. 3. Fernandes VR, Polak JF, Edvardsen T, Carvalho B, Gomes A, et al. Subclinical atherosclerosis and incipient regional myocardial dysfunction in asymptomatic individuals: the multi-ethnic study of atherosclerosis (MESA). J Am Coll Cardiol 2006, 47(12):2420–2428. 4. Zheng ZJ, Croft JB, Giles WH, Mensah GA. Sudden cardiac death in the United States, 1989 to 1998. Circulation 2001, 104(18):2158–2163. 5. Naghavi M, Libby P, Falk E, Casscells SW, Litovsky S, et al. From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: part I. Circulation 2003, 108(14):1664–1672. 6. Fischer A, Gutstein DE, Fayad ZA, Fuster V. Predicting plaque rupture: enhancing diagnosis and clinical decision-making in coronary artery disease. Vasc Med 2000, 5(3):163–172. 7. Gupta AK, Naregalkar RR, Vaidya VD, Gupta M. Recent advances on surface engineering of magnetic iron oxide nanoparticles and their biomedical applications. Nanomedicine 2007, 2(1):23–39. 8. Wickline SA, Neubauer AM, Winter PM, Caruthers SD, Lanza GM. Molecular imaging and therapy of

318

atherosclerosis with targeted nanoparticles. J Magn Reson Imaging 2007, 25(4):667–680. 9. Hyafil F, Cornily JC, Feig JE, Gordon R, Vucic E, et al. Noninvasive detection of macrophages using a nanoparticulate contrast agent for computed tomography. Nat Med 2007, 13(5):636–641. 10. Kaufmann BA, Lindner JR. Molecular imaging with targeted contrast ultrasound. Curr Opin Biotechnol 2007, 18(1):11–16. 11. Frias JC, Ma Y, Williams KJ, Fayad ZA, Fisher EA. Properties of a versatile nanoparticle platform contrast agent to image and characterize atherosclerotic plaques by magnetic resonance imaging. Nano Lett 2006, 6(10):2220–2224. 12. Nahrendorf M, Jaffer FA, Kelly KA, Sosnovik DE, Aikawa E, et al. Noninvasive vascular cell adhesion molecule–1 imaging identifies inflammatory activation of cells in atherosclerosis. Circulation 2006, 114(14):1504–1511. 13. Caruthers SD, Wickline SA, Lanza GM. Nanotechnological applications in medicine. Curr Opin Biotechnol 2007, 18(1):26–30. 14. Emerich DF, Thanos CG. Targeted nanoparticlebased drug delivery and diagnosis. J Drug Target 2007, 15(3):163–183. 15. Wagner V, Dullaart A, Bock AK, Zweck A. The emerging nanomedicine landscape. Nat Biotechnol 2006, 24(10):1211–1217.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

16. Buxton D. The promise of nanotechnology for heart, lung and blood diseases. Expert Opin Drug Deliv 2006, 3(2):173–175. 17. Cyrus T, Winter PM, Caruthers SD, Wickline SA, Lanza GM. Magnetic resonance nanoparticles for cardiovascular molecular imaging and therapy. Expert Rev Cardiovasc Ther 2005, 3(4):705–715. 18. Lanza G, Winter P, Cyrus T, Caruthers S, Marsh J, et al. Nanomedicine opportunities in cardiology. Ann N Y Acad Sci 2006, 1080:451–465. 19. Glagov S, Weisenberg E, Zarins CK, Stankunavicius R, Kolettis GJ. Compensatory enlargement of human atherosclerotic coronary arteries. N Engl J Med 1987, 316(22):1371–1375. 20. Arbustini E, Dal Bello B, Morbini P, Burke AP, Bocciarelli M, et al. Plaque erosion is a major substrate for coronary thrombosis in acute myocardial infarction. Heart 1999, 82(3):269–272. 21. Burke AP, Farb A, Malcom GT, Liang YH, Smialek J, et al. Coronary risk factors and plaque morphology in men with coronary disease who died suddenly. N Engl J Med 1997, 336(18):1276–1282. 22. Virmani R, Kolodgie FD, Burke AP, Farb A, Schwartz SM. Lessons from sudden coronary death: a comprehensive morphological classification scheme for atherosclerotic lesions. Arterioscler Thromb Vasc Biol 2000, 20(5):1262–1275. 23. Davies MJ. Stability and instability: two faces of coronary atherosclerosis. The Paul Dudley White lecture 1995. Circulation 1996, 94(8):2013–2020. 24. Eliceiri BP, Cheresh DA. Role of alpha v integrins during angiogenesis. Cancer J 2000, 6(suppl 3):S245–S249. 25. Kerr JS, Mousa SA, Slee AM. Alpha(v)beta(3) integrin in angiogenesis and restenosis. Drug News Perspect 2001, 14(3):143–150.

human breast cancer growth and angiogenesis in human skin. J Clin Invest 1995, 96(4):1815–1822. 31. Corjay MH, Diamond SM, Schlingmann KL, Gibbs SK, Stoltenborg JK, et al. Alphavbeta3, alphavbeta5, and osteopontin are coordinately upregulated at early time points in a rabbit model of neointima formation. J Cell Biochem 1999, 75(3):492–504. 32. Brooks PC, Montgomery AM, Rosenfeld M, Reisfeld RA, Hu T, et al. Integrin alpha v beta 3 antagonists promote tumor regression by inducing apoptosis of angiogenic blood vessels. Cell 1994, 79(7):1157–1164. 33. Kim S, Bell K, Mousa SA, Varner JA. Regulation of angiogenesis in vivo by ligation of integrin alpha5beta1 with the central cell-binding domain of fibronectin. Am J Pathol 2000, 156(4):1345–1362. 34. Ferrara N, Gerber HP, LeCouter J. The biology of VEGF and its receptors. Nat Med 2003, 9(6):669–676. 35. Weissleder R. Molecular imaging: exploring the next frontier. Radiology 1999, 212(3):609–614. 36. Caruthers SD, Wickline SA, Lanza GM. Targeted nanoparticles for molecular imaging and therapy: a multi-modality approach to molecular medicine. In: Spekowius G, Wendler T: eds. Advances in Healthcare Technology, Vol. 6. Dordrecht: Springer; 2006, 305–322. 37. Lanza G, Wallace K, Scott M, Cacheris W, Abendschein D, et al. A novel site-targeted ultrasonic contrast agent with broad biomedical application. Circulation 1996, 94:3334–3340. 38. Lane TA. Perfluorochemical-based artificial oxygen carrying red cell substitutes. Transfus Sci 1995, 16(1):19–31.

26. Jain RK, Finn AV, Kolodgie FD, Gold HK, Virmani R. Antiangiogenic therapy for normalization of atherosclerotic plaque vasculature: a potential strategy for plaque stabilization. Nat Clin Pract Cardiovasc Med 2007, 4(9):491–502.

39. Croce MA, Fabian TC, Patton JH Jr, Melton SM, Moore M, et al. Partial liquid ventilation decreases the inflammatory response in the alveolar environment of trauma patients. J Trauma 1998, 45(2):273–280, discussion 280–272.

27. Ruehm SG, Corot C, Vogt P, Cristina H, Debatin JF. Ultrasmall superparamagnetic iron oxide-enhanced MR imaging of atherosclerotic plaque in hyperlipidemic rabbits. Acad Radiol 2002, 9(suppl 1):S143–S144.

40. Hirschl RB, Croce M, Gore D, Wiedemann H, Davis K, et al. Prospective, randomized, controlled pilot study of partial liquid ventilation in adult acute respiratory distress syndrome. Am J Respir Crit Care Med 2002, 165(6):781–787.

28. Mulder WJM, Strijkers GJ, Briley-Saboe KC, Frias JC, Aguinaldo JGS, et al. Molecular imaging of macrophages in antherosclerotic plaques using bimodal PEG-micelles. Magn Reson Med 2007, 58(6):1164–1170.

41. Wootton SL, Coley BD, Hilton SV, Edwards DK III, Amberg JR, et al. Value of brominated fluorocarbons for the radiographic diagnosis of small-bowel obstruction: comparison with other contrast agents in rats. Am J Roentgenol 1993, 161(2):409–416.

29. Goldstein JA. Multifocal coronary plaque instability. Prog Cardiovasc Dis 2002, 44(6):449–454.

42. Brown JJ, Duncan JR, Heiken JP, Balfe DM, Corr AP, et al. Perfluoroctylbromide as a gastrointestinal contrast agent for MR imaging: use with and without glucagon. Radiology 1991, 181(2):455–460.

30. Brooks PC, Stromblad S, Klemke R, Visscher D, Sarkar FH, et al. Antiintegrin alpha v beta 3 blocks

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

319

Advanced Review

www.wiley.com/wires/nanomed

43. Clark LC Jr, Clark EW, Moore RE, Kinnett DG, Inscho EI Jr. Room temperature-stable biocompatible fluorocarbon emulsions. Prog Clin Biol Res 1983, 122:169–180. 44. Police AM, Waxman K, Tominaga G. Pulmonary complications after fluosol administration to patients with life-threatening blood loss. Crit Care Med 1985, 13(2):96–98. 45. Flaim SF. Pharmacokinetics and side effects of perfluorocarbon-based blood substitutes. Artif Cells Blood Substit Immobil Biotechnol 1994, 22(4):1043–1054. 46. McGoron AJ, Pratt R, Zhang J, Shiferaw Y, Thomas S, et al. Perfluorocarbon distribution to liver, lung and spleen of emulsions of perfluorotributylamine (FTBA) in pigs and rats and perfluorooctyl bromide (PFOB) in rats and dogs by 19F NMR spectroscopy. Artif Cells Blood Substit Immobil Biotechnol 1994, 22(4):1243–1250. 47. Keipert PE, Otto S, Flaim SF, Weers JG, Schutt EA, et al. Influence of perflubron emulsion particle size on blood half-life and febrile response in rats. Artif Cells Blood Substit Immobil Biotechnol 1994, 22(4):1169–1174. 48. Krafft M. Fluorocarbons and fluorinated amphiphiles in drug delivery and biomedical research. Adv Drug Deliv Rev 2001, 47:209–228. 49. Kobayashi H, Brechbiel MW. Dendrimer-based macromolecular MRI contrast agents: characteristics and application. Mol Imaging 2003, 2(1):1–10. 50. Li KC, Bednarski MD. Vascular-targeted molecular imaging using functionalized polymerized vesicles. J Magn Reson Imaging 2002, 16(4):388–393. 51. Moffat BA, Reddy GR, McConville P, Hall DE, Chenevert TL, et al. A novel polyacrylamide magnetic nanoparticle contrast agent for molecular imaging using MRI. Mol Imaging 2003, 2(4):324–332. 52. Quintana A, Raczka E, Piehler L, Lee I, Myc A, et al. Design and function of a dendrimer-based therapeutic nanodevice targeted to tumor cells through the folate receptor. Pharm Res 2002, 19(9):1310–1316. 53. Sipkins DA, Cheresh DA, Kazemi MR, Nevin LM, Bednarski MD, et al. Detection of tumor angiogenesis in vivo by alphaVbeta3-targeted magnetic resonance imaging. Nat Med 1998, 4:623–626. 54. Unger EC, Porter T, Culp W, Labell R, Matsunaga T, et al. Therapeutic applications of lipidcoated microbubbles. Adv Drug Deliv Rev 2004, 56(9):1291–1314. 55. O’Reilly RK, Hawker CJ, Wooley KL. Cross-linked block copolymer micelles: functional nanostructures of great potential and versatility. Chem Soc Rev 2006, 35(11):1068–1083.

320

56. Majoros IJ, Myc A, Thomas T, Mehta CB, Baker JR Jr. PAMAM dendrimer-based multifunctional conjugate for cancer therapy: synthesis, characterization, and functionality. Biomacromolecules 2006, 7(2):572–579. 57. Lanza GM, Lamerichs R, Caruthers S, Wickline SA. Molecular imaging in MR with targeted paramagnetic nanoparticles. Medicamundi 2003, 47(1):34–39. 58. Paganelli G, Magnani P, Fazio F. Pretargeting of carcinomas with the avidin-biotin system. Int J Biol Markers 1993, 8(3):155–159. 59. Diamandis EP, Christopoulos TK. The biotin(strept)avidin system: principles and applications in biotechnology. Clin Chem 1991, 37(5):625–636. 60. Rosebrough SF. Two-step immunological approaches for imaging and therapy. Q J Nucl Med 1996, 40(3):234–251. 61. Tymkewycz PM, Creighton Kempsford LJ, Gaffney PJ. Generation and partial characterization of five monoclonal antibodies with high affinities for fibrin. Blood Coagul Fibrinolysis 1993, 4(2):211–221. 62. Gaffney PJ, Gascoine PS, Creighton LJ, Tymkewycz PM. Monoclonal antibodies for the detection of thrombosis. Adv Exp Med Biol 1990, 281:419–427. 63. Lanza GM, Trousil RL, Wallace KD, Rose JH, Hall CS, et al. In vitro characterization of a novel, tissue-targeted ultrasonic contrast system with acoustic microscopy. J Acoust Soc Am 1998, 104(6):3665–3672. 64. Hall CS, Marsh JN, Scott MJ, Gaffney PJ, Wickline SA, et al. Temperature dependence of ultrasonic enhancement with a site-targeted contrast agent. J Acoust Soc Am 2001, 110(3 Pt 1):1677–1684. 65. Marsh JN, Hall CS, Wickline SA, Lanza GM. Temperature dependence of acoustic impedance for specific fluorocarbon liquids. J Acoust Soc Am 2002, 112(6):2858–2862. 66. Marsh JN, Hall CS, Scott MJ, Fuhrhop RW, Gaffney PJ, et al. Improvements in the ultrasonic contrast of targeted perfluorocarbon nanoparticles using an acoustic transmission line model. IEEE Trans Ultrason Ferroelectr Freq Control 2002, 49(1):29–38. 67. Marsh JN, Partlow KC, Abendschein DR, Scott MJ, Lanza GM, et al. Molecular imaging with targeted perfluorocarbon nanoparticles: quantification of the concentration dependence of contrast enhancement for binding to sparse cellular epitopes. Ultrasound Med Biol 2007, 33(6):950–958. 68. Hughes MS, Marsh JN, Hall CS, Fuhrhop RW, Lacy EK, et al. Acoustic characterization in whole blood and plasma of site-targeted nanoparticle ultrasound contrast agent for molecular imaging. J Acoust Soc Am 2005, 117(2):964–972. 69. Hughes MS, Marsh JN, Hall CS, Savery D, Lanza GM, et al. Characterization of digital waveforms using

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

thermodynamic analogs: applications to detection of materials defects. IEEE Trans Ultrason Ferroelectr Freq Control 2005, 52(9):1555–1564. 70. Hughes MS, McCarthy JE, Marsh JN, Arbeit JM, Neumann RG, et al. Properties of an entropybased signal receiver with an application to ultrasonic molecular imaging. J Acoust Soc Am 2007, 121(6):3542–3557. 71. Lanza G, Lorenz C, Fischer S, Scott M, Cacheris W, et al. Enhanced detection of thrombi with a novel fibrin-targeted magnetic resonance imaging agent. Acad Radiol 1998, 5(suppl 1):S173–S176. 72. Winter PM, Caruthers SD, Yu X, Song SK, Chen J, et al. Improved molecular imaging contrast agent for detection of human thrombus. Magn Reson Med 2003, 50(2):411–416. 73. Morawski AM, Winter PM, Crowder KC, Caruthers SD, Fuhrhop RW, et al. Targeted nanoparticles for quantitative imaging of sparse molecular epitopes with MRI. Magn Reson Med 2004, 51(3):480–486. 74. Yu X, Song SK, Chen J, Scott MJ, Fuhrhop RJ, et al. High-resolution MRI characterization of human thrombus using a novel fibrin-targeted paramagnetic nanoparticle contrast agent. Magn Reson Med 2000, 44(6):867–872.

magnetic resonance imaging. Nat Clin Pract Cardiovasc Med 2004, 1(1):48–55. 82. Ala-Korpela M, Sipola P, Kaski K. Characterization and molecular detection of atherothrombosis by magnetic resonance–potential tools for individual risk assessment and diagnostics. Ann Med 2006, 38(5):322–336. 83. Moulton K. Plaque angiogenesis and atherosclerosis. Curr Atheroscler Rep 2001, 3:225–233. 84. Lendon C, Born GV, Davies MJ, Richardson PD. Plaque fissure: the link between atherosclerosis and thrombosis. Nouv Rev Fr Hematol 1992, 34(1):27–29. 85. Look DC, Locker DR. Time saving in measurement of NMR and EPR relaxation times. Rev Sci Instrum 1970, 41(2):621–627. 86. Hinshaw WS, Lent AH. An introduction to NMR imaging: from the Bloch equation to the imaging equation. Proc IEEE 1983, 71(3):338–354. 87. Haake EM, Brown RW, Thompson MR, Venkatesan R. Magnetic Resonance Imaging: Physical Principles and Sequence Design. New York: John Wiley & Sons; 1999:, 914.

75. Flacke S, Fischer S, Scott M, Fuhrhop R, Allen J, et al. A novel MRI contrast agent for molecular imaging of fibrin: implications for detecting vulnerable plaques. Circulation 2001, 104:1280–1285.

88. Fan X, River J, Zamora M, Al-Hallaq H, Karczmar G. Effect of carbogen on tumor oxygenation: combined fluorine-19 and proton MRI measurements. Int J Radiat Oncol Biol Phys 2002, 54:1202–1209.

76. Anderson SA, Rader RK, Westlin WF, Null C, Jackson D, et al. Magnetic resonance contrast enhancement of neovasculature with alpha(v)beta(3)targeted nanoparticles. Magn Reson Med 2000, 44(3):433–439.

89. Hunjan S, Zhao D, Canstandtinescu A, Hahan E, Antich P, et al. Tumor oximetry: demonstration of an enhanced dynamic mapping procedure using fluorine19 echo planar magnetic resonance imaging the Dunning prostate R3327-At1 rat tumor. Int J Radiat Oncol Biol Phys 2001, 49:1097–1108.

77. Winter PM, Morawski AM, Caruthers SD, Fuhrhop RW, Zhang H, et al. Molecular imaging of angiogenesis in early-stage atherosclerosis with alpha(v)beta3integrin-targeted nanoparticles. Circulation 2003, 108(18):2270–2274. 78. Zhang Y, Cliff WJ, Schoefl GI, Higgins G. Immunohistochemical study of intimal microvessels in coronary atherosclerosis. Am J Pathol 1993, 143(1):164–172. 79. Moulton KS. Plaque angiogenesis: its functions and regulation. Cold Spring Harb Symp Quant Biol 2002, 67:471–482.

90. Noth U, Grohn P, Jork A, Zimmermann U, Haase A, et al. 19F-MRI in vivo determination of the partial oxygen pressure in perfluorocarbon-loaded alginate capsules implanted into the peritoneal cavity and different tissues. Magn Reson Med 1999, 42:1039–1047. 91. Ikehira H, Girard F, Obata T, Ito H, Yoshitomi H, et al. A preliminary study for clinical pharmacokinetics of oral fluorine anticancer medicines using the commercial MRI system 19F-MRS. Br J Radiol 1999, 72:584–589.

80. Morawski AM, Winter PM, Caruthers SD, Williams TA, Allen, et al. A semi-automated algorithm for quantification of vessel wall angiogenesis associated with early atherosclerosis using magnetic resonance imaging. Proceedings of the 25th Annual Conference of the IEEE EMBS, 2003, Cancun, Mexico, 743–746.

92. Schlemmer H, Becker M, Bachert P, Dietz A, Rudat V, et al. Alterations of intratumoral pharacokinetics of 5-fluorouracil in head and neck carcinoma during simultaneous radiochemotherapy. Cancer Res 1999, 59:2363–2369.

81. Lipinski MJ, Fuster V, Fisher EA, Fayad ZA. Technology insight: targeting of biological molecules for evaluation of high-risk atherosclerotic plaques with

93. Wolf W, Presant C, Waluch V. 19F-MRS studies of fluorinated drugs in humans. Adv Drug Deliv Rev 2000, 41:55–74.

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

321

Advanced Review

www.wiley.com/wires/nanomed

94. Huang M, Ye Q, Williams D, Ho C. MRI of lungs using partial liquid ventilation with water-inperfluorocarbon emulsions. Magn Reson Med 2002, 48:487–492.

106. Lanza GM, Winter PM, Caruthers SD, Morawski AM, Schmieder AH, et al. Magnetic resonance molecular imaging with nanoparticles. J Nucl Cardiol 2004, 11(6):733–743.

95. Laukemper-Ostendorf S, Scholz A, Burger K, Heussel C, Schmittner M, et al. 19F-MRI of perflubron for measurement of oxygen partial pressure in porcine lungs during partial liquid ventilation. Magn Reson Med 2002, 47:82–89.

107. Leary SP, Liu CY, Apuzzo ML. Toward the emergence of nanoneurosurgery: part III–nanomedicine: targeted nanotherapy, nanosurgery, and progress toward the realization of nanoneurosurgery. Neurosurgery 2006, 58(6):1009–1026, discussion 1009–1026.

96. Morawski AM, Winter PM, Yu X, Fuhrhop RW, Scott MJ, et al. Quantitative ‘‘magnetic resonance immunohistochemistry’’ with ligand-targeted (19)F nanoparticles. Magn Reson Med 2004, 52(6):1255–1262. 97. Robertson JD, Crane SB, Wickline SA, Lanza GM. Characterization and biodistribution of a novel MRI molecular imaging agent by neutron activation analysis. J Radioanal Nucl Chem 2005, 263(2):511–514. 98. Caruthers SD, Neubauer AM, Hockett FD, Lamerichs R, Winter PM, et al. In vitro demonstration using 19F magnetic resonance to augment molecular imaging with paramagnetic perfluorocarbon nanoparticles at 1.5 Tesla. Invest Radiol 2006, 41(3):305–312. 99. Pettigrew RI, Fee CA, Li KC. Changes in the world of biomedical research are moving the field of ‘‘personalized medicine’’ from concept to reality. J Nucl Med 2004, 45(9):1427. 100. Lanza GM, Yu X, Winter PM, Abendschein DR, Karukstis KK, et al. Targeted antiproliferative drug delivery to vascular smooth muscle cells with a magnetic resonance imaging nanoparticle contrast agent: implications for rational therapy of restenosis. Circulation 2002, 106(22):2842–2847. 101. Wickline SA, Lanza GM. Molecular imaging, targeted therapeutics, and nanoscience. J Cell Biochem Suppl 2002, 39:90–97. 102. Winter PM, Caruthers SD, Wickline SA, Lanza GM. Molecular imaging and targeted drug delivery in cardiovascular disease with paramagnetic nanoparticles. Mol Imaging 2004, 3(3):188. 103. Moulton KS, Vakili K, Zurakowski D, Soliman M, Butterfield C, et al. Inhibition of plaque neovascularization reduces macrophage accumulation and progression of advanced atherosclerosis. Proc Natl Acad Sci U S A 2003, 100(8):4736–4741. 104. Moulton KS, Heller E, Konerding MA, Flynn E, Palinski W, et al. Angiogenesis inhibitors endostatin or TNP-470 reduce intimal neovascularization and plaque growth in apolipoprotein E-deficient mice. Circulation 1999, 99(13):1726–1732. 105. Ingber D, Fujita T, Kishimoto S, Sudo K, Kanamaru T, et al. Synthetic analogues of fumagillin that inhibit angiogenesis and suppress tumour growth. Nature 1990, 348(6301):555–557.

322

108. Tomalia DA, Reyna LA, Svenson S. Dendrimers as multi-purpose nanodevices for oncology drug delivery and diagnostic imaging. Biochem Soc Trans 2007, 35(Pt 1):61–67. 109. Crowder KC, Hughes MS, Marsh JN, Barbieri AM, Fuhrhop RW, et al. Sonic activation of molecularlytargeted nanoparticles accelerates transmembrane lipid delivery to cancer cells through contactmediated mechanisms: implications for enhanced local drug delivery. Ultrasound Med Biol 2005, 31(12):1693–1700. 110. Hernandez-Borrell J, Mas F, Puy J. A theoretical approach to describe monolayer-liposome lipid interaction. Biophys Chem 1990, 36(1):47–55. 111. Jahnig F. Lipid exchange between membranes. Biophys J 1984, 46(6):687–694. 112. Jones JD, Thompson TE. Mechanism of spontaneous, concentration-dependent phospholipid transfer between bilayers. Biochemistry 1990, 29(6):1593–1600. 113. Winter PM, Morawski AM, Caruthers SD, Harris TD, Fuhrhop RW, et al. Antiangiogenic therapy of early atherosclerosis with paramagnetic alpha(v)beta(3)-integrin-targeted fumagillin nanoparticles. J Am Coll Cardiol 2004, 43(5):322A–323A. 114. Winter PM, Neubauer AM, Caruthers SD, Harris TD, Robertson JD, et al. Endothelial alpha(v)beta3 integrin-targeted fumagillin nanoparticles inhibit angiogenesis in atherosclerosis. Arterioscler Thromb Vasc Biol 2006, 26(9):2103–2109. 115. Winter PM, Morawski AM, Caruthers SD, Harris TD, Hu G, et al. Serial quantification of targeted fumagillin therapy using avb3-targeted paramagnetic nanoparticles in early atherosclerosis at 1.5T. J Cardiovasc Magn Reson 2005, 7(1):3–3. 116. Winter PM, Morawski AM, Caruthers SD, Harris TD, Hu G, et al. Persistent antiangiogenic therapy with integrin-targeted fumagillin nanoparticles in early atherosclerosis. J Am Coll Cardiol 2005, 45(3):435A–435A. 117. Winter PM, Caruthers SD, Allen JS, Williams TA, Zhang HY, et al. Combination therapy of targeted anti-angiogenic drug delivery and oral statin against atherosclerosis. J Am Coll Cardiol 2007, 499:151A–151A.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Anti-angiogenic perfluorocarbon nanoparticles for diagnosis and treatment of atherosclerosis

FURTHER READING American Heart Association. ‘‘Atherosclerosis.’’ http://www.americanheart.org/presenter.jhtml?identifier=4440 Associationfor the Eradication of Heart Attacks. ‘‘The Vulnerable Plaque’’ Multimedia. http://www.aeha.org/ Multimedia%20-%20Association%20for%20Eradication%20of%20Heart%20Attacks.htm NationalInstitutes of Health. ‘‘Roadmap for Medical Research: Nanomedicine.’’ http://nihroadmap. nih.gov/nanomedicine/

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

323

Advanced Review

Cell-targeting and cell-penetrating peptides for delivery of therapeutic and imaging agents Rudolph L Juliano,∗ Rowshon Alam,1 Vidula Dixit1 and Hyun Min Kang1 This review will discuss the basic concepts concerning the use of cell-targeting peptides (CTPs) and cell-penetrating peptides (CPPs) in the context of nanocarrier technology. It deals with the discovery and subsequent evolution of CTPs and CPPs, issues concerning their interactions with cells and their biodistribution in vivo, and their potential advantages and disadvantages as delivery agents. The article also briefly discusses several specific examples of the use of CTPs or CPPs to assist in the delivery of nanoparticles, liposomes, and other nanocarriers.  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 324–335

BASIC CONCEPTS

T

he development of effective nanocarrier delivery systems for therapeutic and imaging agents often requires means for highly selective targeting of the nanocarrier to particular cells or tissues. Additionally, in many cases, entry of the nanocarrier into the cell interior is also required. Therefore much effort has recently been devoted to devising means for promoting selective uptake of nanocarriers by cells. Two of the most promising strategies in this arena have been based on the use of relatively small peptides to promote selective binding and effective internalization. Thus ‘cell-targeting peptides’ (CTPs) connote a diverse group of molecules that have emerged from library screening, or by design, to bind to specific cell-surface receptors with high affinity and selectivity. Binding may or may not promote efficient cell entry, and therefore another group of peptides have been developed, the ‘cell-penetrating peptides’ (CPPs), that can enhance passage of molecules or nanoparticles across membrane barriers. Use of CTPs and CPPs individually or jointly can provide us with powerful tools to enhance the capabilities and utility of nanocarriers. In the sections below, we will discuss a number of aspects of both CTPs and CPPs as applied to nanomedicine. This discussion complements several excellent recent

∗

Correspondence to: [email protected]

1 Department

of Pharmacology, University of North Carolina, Chapel Hill, NC, USA DOI: 10.1002/wnan.004

324

articles that have dealt with the overall issues regarding targeting and biodistribution of nanoparticles and polymers and their potential use in therapeutics,1–9 and others that have focused explicitly on the chemistry and biology of CPPs10 and CTPs.11–13

SOURCES OF CELL-TARGETING PEPTIDES CTPs have arisen from many different sources. One obvious possibility is to utilize peptides that are natural ligands for important cell-surface receptors. A good example of this is transferrin, an iron-binding polypeptide that interacts with high affinity with its cognate receptor and is then carried into the cell via receptor-mediated endocytosis. Transferrin has been used to decorate a variety of nanoparticle and macromolecular carriers and has been effective in promoting cellular uptake in culture and in vivo.14 Transferrin is thought to be expressed at higher levels in rapidly growing tumor cells than in normal cells,15 and therefore it has been popular in terms of tumor-directed targeting. However, most cells express some level of transferrin receptor, and therefore the selectivity attained is only partial. In a similar vein, peptide ligands for growth-factor receptors that are important in cancer, such as EGF-R, have also been used as CTPs.16–18 Since natural growth factors are fairly large molecules, in some cases shorter peptides or peptidomimetics have served as ligands for growth-factor receptors, and some of these have been used for delivery purposes.19,20

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

An important class of CTPs has emerged from studies of the ligand-binding properties of members of the integrin family of cell-adhesion receptors.21 Certain extracellular proteins such as vitronectin and fibronectin interact with their cognate integrins primarly via the tripeptide RGD (Arg-Gly-Asp). However, the affinity of RGD-binding integrins for a simple RGD sequence is very poor. Therefore a great deal of effort has gone into the development of modified peptides or peptidomimetics having higher affinity as well as good selectivity for particular integrins including α4 β1 , αIIb β3 , and αv β3 22–24 This effort has proceeded furthest in the context of αIIb β3 , the fibrinogen-binding platelet integrin. Thus orally active compounds with nanomolar affinities for αII β3 are in the clinic, and are used for control of blood-clotting in various contexts, although these drugs are not without problems.25 The αv β3 integrin is highly involved in the process of angiogenesis and is also preferentially expressed on certain types of tumor cells.26 Thus there has been a substantial interest in developing high-affinity ligands for αv β3 both as direct therapeutic agents as well as for targeting. Cyclization and modification of the RGD moiety has led to many compounds with improved affinity. Recently, several laboratories have used bivalent or multivalent cyclic RGD moieties to attain quite high affinity (∼10 nM) for αv β3 ,27,28 while nonpeptide organic molecules have also been developed and are in clinical trials for various cancers.29,30 Cyclic RGD peptides have been coupled to agents used in various imaging modalities (magnetic resonance imaging (MRI), positron emission tomography (PET), fluorescence) and have been very effective in visualizing αv β3 −containing tumors and tissues in living animals.31 The most general and most powerful strategy for generating novel CTPs is to use combinatorial library approaches such as phage display, RNA display, and bacterial or yeast display.32–35 In simplest form, a particular cell-surface receptor is expressed and purified and is then used as a substrate for screening a large peptide library for high-affinity ligands. Often, multiple rounds of screening followed by optimization of the leads obtained from the library are needed to attain robust and highly selective ligand–receptor binding. An important variation on this process is ‘in vivo phage display’.36 Thus phages expressing a peptide library are injected into animals (mice) and, after a period to allow clearance of nonbinding particles, the residual phages are harvested from various tissues or tumors. Remarkably, after several such cycles, it is possible to harvest phages that exhibit specificity for particular cell and tissue types. For example, this strategy has led to phages Vo lu me 1, May /Ju n e 2009

that target tumor lymph channels,37 organ-specific vascular endothelium,38 or tumor determinants.39 Initially, the cell-surface receptor involved in tissuespecific binding is not known; however, in several cases the responsible protein has been identified.40,41 While the in vivo phage display technique is potentially very powerful in identifying new cell-type-specific receptors and cognate peptide ligands, there are some limitations. Firstly, because of the relatively large size of the phage particles, it is mainly the intravascular space that is probed.42 Therefore most of the receptors identified by this approach have been on endothelial cells. Second, although the phage may bind to a particular cell type with high affinity, this is often due to the fact that several copies of the library peptide are displayed on the phage surface. When the peptides are tested in monovalent form, they sometimes have rather low affinities, and thereby prompting a great deal of additional work on optimization. There is a continuum of cell-targeting agents in terms of size and complexity. Most of the peptides identified through phage display or other combinatorial library approaches are relatively small molecules, perhaps with 6–12 amino acids. Natural ligands for cell surface receptors are considerably larger; for example, transferrin is about 80 kDa,43 while EGF is smaller at 53 amino acids.44 At the other end of the scale, monoclonal antibodies provide superb cell-type-specific targeting capabilities, but they are rather large proteins with molecular weights in the 160 kDa range and are subject to denaturation because of their complex structures. An intermediate situation relates to engineered fragments or variants of antibodies such as the classic Fab fragments, scFvs that combine light- and heavychain variable regions in a single polypeptide, and bivalent, trivalent, and tetravalent ‘minibodies’.32,45 Other intermediate-sized reagents can be derived from peptide libraries inserted into small, stable protein domains, such as the fibronectin type III repeat domain,46 or from natural full proteome libraries.47 Thus a plethora of cell-targeting reagents having various characteristics in terms of size, stability, affinity, and specificity are available to apply to the issue of delivery of nanocarriers. Appropriate choice of a targeting agent will depend on the problem at hand, and no one approach will be suited for every situation.

 2009 Jo h n Wiley & So n s, In c.

325

Advanced Review

www.wiley.com/wires/nanomed

CHEMISTRY AND BIOLOGY OF CELL-PENETRATING PEPTIDES During the last decade, an important, new approach to the intracellular delivery of macromolecules and nanocarriers has emerged. This is based on the socalled ‘protein-transduction domains’ (PTDs) also known as CPPs. The prototypical CPPs are short cationic peptides (TAT, ANT) derived from the transcriptional regulator proteins HIV Tat and drosophila Antennepedia;10,48 ‘TAT’ and ‘ANT’ have now been joined by a large number of additional CPPs.49,50 Many CPPs have a polycationic character, but others are based on hydrophobic sequences derived from signal peptides, viral peptides, or other sources.51 CPPs can not only enter cells themselves but, with greater or lesser efficiency, can also transport attached ‘cargo’ molecules. However, the efficiency of delivery is highly affected by the nature of the cargo.52 Certain CPPs have very effectively delivered, biologically active (but membrane impermeant) short peptides, thereby allowing some elegant mechanistic studies of the role of these active peptides in signaling processes.53,54 Cationic and hydrophobic CPPs have also been reported to permit intracellular delivery of proteins into cultured cells,55 as well as in vivo delivery of enzymes such as β-galactosidase and Cre recombinase to tissues.56,57 However, some of the studies on CPP-mediated protein delivery, especially those in vivo, have been controversial, and some investigators have found that current CPPs are rather ineffective in promoting intracellular delivery of proteins.58,59 TAT and ANT variety of CPPs have also been used for the intracellular delivery of antisense and siRNA oligonucleotides,60–64 but once again results have been somewhat controversial.65 Several reports suggest that even the delivery of large entities such as liposomes and magnetic nanoparticles66,67 can be enhanced via CPPs, but the generality of this approach is unclear. Although various CPPs can cause cytotoxicity when used at high levels, for the most part they are relatively nontoxic when used at low concentrations.49 Concepts concerning the mechanism(s) of cell entry by CPPs have been undergoing a rapid evolution. Prior to discussing these studies, it seems appropriate to briefly review aspects of endogenous pathways for trafficking of macromolecules in cells. Many cells utilize multiple mechanisms of endocytosis to allow import of large and small molecules; this includes classic clathrin-mediated endocytosis, smooth vesicle pinocytosis, actin-dependent macropinocytosis, and entry via glycosphingolipid-rich calveolae.68 Interestingly, viruses as well as bacterial and plant toxins exploit many of the cellular pathways of 326

macromolecule importation.69–71 For example, many viruses, including SV40, Semliki Forest Virus, and Influenza A, primarily enter cells via clathrin-coated pits and then use the low-pH environment of early endosomes to trigger a conformational change in viral surface proteins, which then allows escape from the endosome into the cytosol. Early work on prototypical CPPs such as TAT and ANT suggested that they could directly cross the plasma membrane and enter the cytosol.72 However, this interpretation was based on fluorescence microscopy studies with fixed cells, and it has become apparent that the fixation process caused artifacts.73 More recent live-cell studies indicate that most cationic CPPs enter cells by binding to cell-surface proteoglycans,74 followed by uptake into endosomes most likely by macropinocytosis,75,76 followed by partial release from endosomes via a pH-dependent mechanism. As a result of this process, substantial amounts of these cationic peptides (and their cargos) remain within the endosomal compartment. Another important issue, alluded to above, is that the mechanism and extent of cell uptake of CPPs often depend on the nature of the associated ‘cargo’ molecule.77 Not surprisingly, one might expect that a CPP linked to a small peptide might undergo a different cell entry process than CPPs linked to a much larger nanocarrier. The mechanism(s) involved in the passage of CPPs and their cargos across endo-membranes are still poorly understood. Recent biophysical studies suggest that cationic CPPs interact with anionic phospholipid head groups to trigger the formation of destabilizing nonbilayer domains in the membrane.78,79 However, the permeation process involving CPPs does not involve a generalized disruption of plasma or endosomal membranes, since several studies have shown that there is no escape of cytosolic marker enzymes.49 Many of the mechanistic studies of CPP actions have involved the TAT and ANT peptides or closely related molecules; in contrast, many of the newer examples of this class of molecules have received much less mechanistic scrutiny and therefore little is known about how they enter cells.

ISSUES REGARDING IN VIVO DELIVERY AND TOXICITY An important issue for the use of CPP (or CTPs) for nanocarrier delivery is the biodistribution of the peptides themselves or when attached to protein or nucleic acid cargos; this has been explored only to a very limited degree. In one study, the pharmacokinetics and organ distribution of free TAT peptide were compared to those of

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

TABLE 1 Peptide–protein Conjugates CPP or CTP

Nanocarrier

Therapeutic or imaging agent

In vivo or in vitro

Purpose of study

RGD

PEG–albumin

P38 MAPK inhibitor

In vitro

Inhibition of angiogenesis

84

RGD

PEG–albumin

Auristatin (anti-cancer drug)

In vitro

Targeting of tumor or endothelial cells

85

Arginine-rich cyclic peptides

Albumin

NA

In vitro

Screen for cell-penetrating peptides

86

Arginine-rich peptides

NA

Insulin

In vivo

Enteric delivery of insulin

87

Reference

TABLE 2 Peptide–nanoparticle Conjugates CTP or CPP

Nanocarrier

Therapeutic or imaging agent

Purpose of study

Reference

RGD

3-Aminopropyltrimethoxysilane (APTMS)

Iron oxide

In vivo MRI

88

Thiolated peptidomimetic vitronectin antagonist

N-[{w-[4-(p-maleimidophenyl) butanoyl]amino} poly(ethylene glycol)2000 ] 1,2-distearoyl-snglycero-3-phosphoethanolamine (MPB-PEG-DSPE)

Gadolinium (Gd3+ )

In vivo MRI

89

TAT

Aminated dextran

Iron oxide

In vivo MRI, cell tracking

90

TAT

Aminated dextran

Iron oxide, fluorochromes (VT-680, AF680, Cy5 and Cy5.5)

In vivo imaging via fluorescence and MRI

91

Transferrin

PGLA

Paclitaxel

In vivo tumor therapy

92

Transferrin

Cyclodextrin

siRNA

In vivo preclinical toxicology of siRNA in a nanoparticle

93

Deslorelin, transferrin

Polystyrene nanoparticle

Ex vivo

94

Transferrin

Mercaptoundecanoic acid, lysine

In vitro imaging of cancer cells

95

the TAT peptide conjugated to streptavidin. The free peptide was cleared very rapidly, whereas the conjugate had a longer circulation lifetime but less than that of the unmodified protein; only modest effects were observed on uptake into individual organs.80 A very different picture was presented in a study in which TAT or ANT CPPs were coadministered with an antitumor single-chain antibody; here the presence of the CPPs substantially enhanced retention of the antibody by the tumor.81 In contrast, a study using conjugates of TAT peptides with tumor-targeting antibody fragments observed reduced tumor uptake of the conjugates versus the unmodified antibodies.82 Another study undertook Vo lu me 1, May /Ju n e 2009

CdSe/CdS/ZnS quantum rods

an examination of the biodistribution of a TATgalactosidase chimera when administered by several routes; surprisingly however, no comparison was made to unmodified β-galactosidase.24 In a study of an arginine-rich CPP conjugated to a morpholinooligonucleotide, it was found that the CPP increased the plasma half-life and area under the curve (AUC), and increased tissue uptake, particularly in liver and spleen.83 This study also provided a preliminary toxicity evaluation of this type of compound, indicating some degree of renal toxicity of the CPP–oligonucleotide conjugate as compared to the oligonucleotide itself, especially at higher doses. These highly divergent results make it clear that it is still very

 2009 Jo h n Wiley & So n s, In c.

327

Advanced Review

www.wiley.com/wires/nanomed

TABLE 3 Peptide–polymer Conjugates CPP or CTP

Nanocarrier

Therapeutic or imaging agent

Purpose of Study

Reference

RGD

PEG–PEI

pCMV-sFit-1

In vivo antiangiogenic tumor therapy

96

CD21 receptor biding peptide (RMWPSSTVNLSAGRR)

HPMA

N.A.

In vitro screening for targeting peptides

97

Mono and doubly cyclized RGD

HPMA

Indium-111

In vivo targeting of avb3 integrin in tumors

98

RGD

PEG–PEI, PEI

pGL3 plasmid

In vitro Gene delivery via integrins

99

Transferrin

PEG–PEI

pCMVL plasmid

In vivo gene delivery to tumors

100

Tat

PEG–PEI

pGL3 plasmid

In vivo DNA delivery to lung

101

CD-13 binding peptide (CNGRC)

PEG–PEI

β-Gal plasmid, YFP plasmid

In vivo gene delivery to tumors

102

Transferrin

PEG–PEI, PEI

CMV fl plasmid

In vivo fluorescence imaging of tissues

103

Tat

HPMA

Dox, FITC, Texas Red

In vitro uptake by tumor cells

104

Tat, Lys9

PEG

siRNA

In vitro siRNA uptake

105

Tat

PLLA–PEG, Poly(methacryloyl sulfadimethoxine)—PEG (PSD–PEG)

Dox

In vitro delivery of antitumor drugs to cells

106

RGD: Arg-Gly-Asp; PEG: Poly(ethyleneglycol); PEI: Polyethyleneimine; HPMA: Hydroxy Poly methacrylate.

early when it comes to assessing the in vivo behavior of chimeras and conjugates of CPPs with proteins or nucleic acids. In addition, there has not been any formal, comprehensive toxicological evaluation of CPPs (or CTPs). Therefore, efforts to link these peptides to nanocarriers are in effect ‘flying blind’ with respect to anticipation of in vivo biodistribution or possible toxicities.

EXAMPLES OF DELIVERY OF NANOCARRIERS USING CELL-TARGETING PEPTIDES AND CELL-PENETRATING PEPTIDES There have been a large number of studies using CTPs or CPPs to enhance delivery of various types of nanocarriers, both in cell culture and in animal models. While too diverse to discuss individually in detail, many of these studies are listed in Tables 1–5. They are grouped according to the nature of the nanocarrier moiety used: for example, liposomes versus polymers versus nanoparticles. In selected cases, aspects of these studies will be discussed. 328

Blood proteins such as albumin have long been used as carriers for drugs. Recently, this approach has been modified to include use of CTPs or CPPs to enhance targeting or uptake of the drug–protein conjugates. A particularly successful example of this involves use of RGD–PEG polymers to target drug–albumin conjugates to angiogenic vasculature.84,85 There is an enormous body of literature on the use of various types of nanoparticles as drug carriers3,5 but relatively few examples that involve use of peptides as adjuncts for delivery. A variety of nanoparticles have been ‘decorated’ with CTPs or CPPs. Particles containing iron or gadolinium as MRI enhancing agents have been surface-modified with RGD derivatives or the TAT CPPs. Transferrin has been used to target polymeric nanoparticles containing conventional anticancer drugs or siRNAs to tumors in vivo with considerable success. There is extensive literature on use of heterodisperse polymers as drug carriers.4,6,8 Recently, there has been considerable activity in connection with using CTPs or CPPs coupled to polymeric systems.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

TABLE 4 Peptide–liposome Conjugates CPP or CTP

Nanocarrier

Therapeutic or imaging agent

Purpose of study

RGD

Sterically stabilized liposome

Doxorubicin (anticancer/ antiproliferative drug)

Improve antitumor efficacy of doxorubicin (in vivo )

107

RGD

Sterically stabilized liposome

Dexamethasone phosphate (anti-inflammatory)

Inhibition of angiogenesis and thereby experimental arthritis (in vivo )

108

RGD

Liposomes

B10 (dodecahydrododecaborate) radiotherapeutic agent for neutron capture therapy

Inhibition of angiogenesis by targeting tumor vasculature (in vitro )

109

RGD

Sterically stabilized liposome

5-Fluorouracil (anticancer agent)

Inhibition of lung metastasis and angiogenesis in mice

110

Transferrin

Sterically stabilized liposome

Citicoline (neuroprotective agent)

Drug targeting to brain by targeting cells of the blood–brain barrier (in vitro )

111

Growth factor antagonist [D-Arg6 , D-Trp7,9 -Nme Phe8 ]substance P(6–11) antagonist G

Sterically stabilized liposome

Doxorubicin (anticancer/antiproliferative drug)

Targeting small-cell lung carcinoma cells (in vitro )

112

Epidermal growth factor

Sterically stabilized liposome

B10 (boronated acridine) radiotherapeutic agent for neutron capture therapy

Boron neutron capture therapy for cancer cells (in vitro )

113

TAT

pH-sensitive PEG-coated liposomes

Nontherapeutic plasmid encoding green fluorescent protein

Development of tumor-specific stimuli-sensitive drug and gene delivery (in vivo )

114

As seen Table 3, a number of studies have targeted polymers bearing drugs, DNA, or imaging agents to the αv β3 integrin via RGD-type peptides. Another popular approach has been to use the TAT peptide to enhance cell or tissue uptake of polymers complexed or conjugated with DNA, siRNA, or the anticancer drug doxorubicin. Liposomes represent one of the oldest nanotechnologies for delivery of drugs and imaging agents.3,7 A great deal has been learned about how to successfully couple targeting agents to the liposome surface while still maintaining membrane integrity. Likewise ‘sterically stabilized’ liposomes were one of the first highly successful applications of PEGylation technology to extend the circulation lifetime of nanocarriers by impeding binding of plasma proteins and uptake by professional phagocytes in liver and spleen. There is an extensive body of literature (not discussed here) on Vo lu me 1, May /Ju n e 2009

Reference

antibody-mediated targeting of liposomes. More germane to the theme of this review, as seen in Table 4, there have been a number of studies using RGD peptides to target tumors and tumor vasculature with antiproliferative drugs. Peptides targeting growth factor receptors or the transferrin receptor have also been used for targeted liposome delivery. CTPs have been directly conjugated to a variety of therapeutic and imaging agents Table 5. An advantage of this approach is that the resulting conjugates are much smaller than nanocarriers and therefore may have better access to many to tissue sites. A disadvantage is that the small size also results in rapid excretion via glomerular filtration. There has been a great deal of interest in using highaffinity RGD-type ligands to target tumors. A variety of imaging agents including radiotracers for PET imaging, gadolinium chelates for MRI enhancement, and near-IR dyes for optical imaging have all been

 2009 Jo h n Wiley & So n s, In c.

329

Advanced Review

www.wiley.com/wires/nanomed

TABLE 5 Direct Conjugates with Drugs and Imaging Agents CPP or CTP

Nanocarrier

Therapeutic or imaging agent

Purpose of study

RGD

PEG

Radiotracer (64 Cu-DOTA-PEG-RGD).

Tumor imaging and therapeutic applications.

Reference 115

RGD

N/A

Doxorubicin-RGD-4C, acts as tumor inhibitor.

Tumor targeting.

116

Dimeric cyclic RGD

N/A

Radiolabeled-RGD peptide, a potential imaging and therapeutic agent.

To study specific tumor uptake as well as therapy of radiolabeled dimeric RGD peptide.

117

Dimeric cyclic RGD

N/A

Paclitaxel, an antimicrotubule agent, a potent antitumor drug.

The potential of tumortargeted delivery of paclitaxel-RGD conjugate and its utilization as antitumor agent.

118

Tetrameric cyclic RGD

N/A

64 Cu-DOTA-E{E[c(RGDfk)]

2 }2 : microPET imaging of glioma integrin αv β3 expression.

To investigate integrin targeting characteristics of 64 Cu-DOTA-E{E[c(RGDfk)] } 2 2 as a potential agent for diagnosis and receptor-mediated internal radiotherapy of integrin receptor-expressing tumors.

119

Multimeric RGD

N/A

Cypate, an optical imaging agent.

Design, synthesis and evaluation of multimeric arrays of RGD peptides on a near-infrared fluorescent dye (cypate) for tumor targeting.

120

RGD-tetramers

N/A

[99m Tc(HYNIC-tetramer) (tricine)(TPPTS)] is a new promising radiotracer for noninvasive imaging of the integrin αv β3 -positive tumors by SPECT.

To explore the impact of peptide multiplicity on biodistribution characteristics and metabolism of the 99m Tc-labeled multimeric cyclic RGDfk peptides.

121

Cyclic RGD

N/A

[99m Tc(CO)3 -cyclo[RGDyk(PZ)]]+ , a potential imaging agent to visualize angiogenesis and tumor formation in vivo.

Targeting integrin receptors upregulated on tumor cells and neovasculature.

122

Neurotensin (NT), a tridecapeptide

N/A

NT-XI, NT-XII, NT-XIII; new NT analogs for imaging tumors.

Development of double-stabilized neurotensin analogs as potential radiopharmaceuticals for the application in tumor imaging and potentially, therapy of NT receptor–positive tumors.

123

Bitistatin (polypeptide)

N/A

Labeled bitistatin, a promising in vivo imaging agent.

Targeting αv β3 integrins in tumor angiogenesis.

124

Transferrin

Transferrin receptor (TfR)

Transferrin (Tf), an anticancer drug delivery agent.

To enhance the intracellular drug release in cultured tumor cells by Tf-oligomers.

125

RGD: Arg-Gly-Asp; PEG: Poly(ethyleneglycol); DOTA: 1,4,7,10-tetraazacyclododecane-N,N’N’’,N’’’-tetraacetic acid; HYNIC: 6hydrazinonicotinamide; TPPTS: trisodium triphenylphosphine-3,3′ , 3′′′ trisulfonate; SPECT: single photon emission computed tomography.

330

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

linked to RGD ligands. Multimerization of the RGD peptides can result in increased affinity of this type of conjugate.

CONCLUSION The basic concept of enhancing the targeting and intracellular delivery of nanocarriers with CTPs and CPPs is both exciting and valid. However, results attained to date, while promising, often fall short of the hoped for improvements in therapy or imaging. We are still largely working with first-generation CTPs and CPPs and there is a need to continue evolution of these technologies. We need CTPs that have higher

affinity and greater specificity for known receptors, and we must define new CTPs for additional targets that may provide greater cell-type selectivity than current reagents such as transferrin and EGF-type ligands. For CPPs, more must be done to investigate the true mechanism of action of these molecules, and to improve the efficiency with which they move from plasma membrane or endomembrane compartments to the cytosol. The issue of how the ‘cargo’ affects the efficiency and mechanism of action of CPPs also needs further clarification. Even with these caveats, however, the potential importance of CTPs and CPPs for controlled and targeted delivery of therapeutic and imaging agents is very clear.

REFERENCES 1. Park TG, Jeong JH, Kim SW. Current status of polymeric gene delivery systems. Adv Drug Deliv Rev 2006, 58(4):467–486.

12. Ruoslahti E. Targeting tumor vasculature with homing peptides from phage display. Semin Cancer Biol 2000, 10(6):435–442.

2. Wagner E. Programmed drug delivery: nanosystems for tumor targeting. Expert Opin Biol Ther 2007, 7(5):587–593.

13. Khandare JJ, Minko T. Antibodies and peptides in cancer therapy. Crit Rev Ther Drug Carrier Syst 2006, 23(5):401–435.

3. Torchilin VP. Multifunctional nanocarriers. Adv Drug Deliv Rev 2006, 58(14):1532–1555.

14. Widera A, Norouziyan F, Shen WC. Mechanisms of TfR-mediated transcytosis and sorting in epithelial cells and applications toward drug delivery. Adv Drug Deliv Rev 2003, 55(11):1439–1466.

4. Kabanov AV. Polymer genomics: an insight into pharmacology and toxicology of nanomedicines. Adv Drug Deliv Rev 2006, 58(15):1597–1621. 5. LaVan DA, McGuire T, Langer R. Small-scale systems for in vivo drug delivery. Nat Biotechnol 2003, 21(10):1184–1191. 6. Lee CC, MacKay JA, Fr´echet JM, Szoka FC. Designing dendrimers for biological applications. Nat Biotechnol 2005, 23(12):1517–1526. 7. Li SD, Huang L. Gene therapy progress and prospects: non-viral gene therapy by systemic delivery. Gene Ther 2006, 13(18):1313–1319. 8. Duncan R. Designing polymer conjugates as lysosomotropic nanomedicines. Biochem Soc Trans 2007, 35(Pt 1):56–60. 9. Majoros IJ, Myc A, Thomas T, Mehta CB, Baker JR Jr. PAMAM dendrimer-based multifunctional conjugate for cancer therapy: synthesis, characterization, and functionality. Biomacromolecules 2006, 7(2):572–579. ¨ 10. Lindgren M, Hallbrink M, Prochiantz A, Langel U. Cell-penetrating peptides. Trends Pharmacol Sci 2000, 21(3):99–103. 11. Sergeeva A, Kolonin MG, Molldrem JJ, Pasqualini R, Arap W. Display technologies: application for the discovery of drug and gene delivery agents. Adv Drug Deliv Rev 2006, 58(15):1622–1654.

Vo lu me 1, May /Ju n e 2009

15. Daniels TR, Delgado T, Rodriguez JA, Helguera G, Penichet ML. The transferrin receptor part I: biology and targeting with cytotoxic antibodies for the treatment of cancer. Clin Immunol 2006, 121(2):144–158. 16. Riese DJ II, Gallo RM, Settleman J. Mutational activation of ErbB family receptor tyrosine kinases: insights into mechanisms of signal transduction and tumorigenesis. Bioessays 2007, 29(6):558–565. 17. Zeng F, Lee H, Allen C. Epidermal growth factorconjugated poly(ethylene glycol)-block- poly(deltavalerolactone) copolymer micelles for targeted delivery of chemotherapeutics. Bioconjug Chem 2006, 17(2):399–409. 18. Blessing T, Kursa M, Holzhauser R, Kircheis R, Wagner E. Different strategies for formation of pegylated EGF-conjugated PEI/DNA complexes for targeted gene delivery. Bioconjug Chem 2001, 12(4):529–537. 19. Li Z, Zhao R, Wu X, Sun Y, Yao M, Li J, Xu Y, Gu J. Identification and characterization of a novel peptide ligand of epidermal growth factor receptor for targeted delivery of therapeutics. FASEB J 2005, 19(14):1978–1985. 20. Li D, Yu H, Huang H, Shen F, Wu X, Li J, Wang J, Cao X, Wang Q, Tang G. FGF Receptor-mediated gene

 2009 Jo h n Wiley & So n s, In c.

331

Advanced Review

www.wiley.com/wires/nanomed

delivery using ligands coupled to polyethylenimine. J Biomater Appl 2007, 22(2):163–180. 21. Hynes RO. Integrins: bidirectional, allosteric signaling machines. Cell 2002, 110(6):673–687. 22. Tilley JW, Chen L, Sidduri A, Fotouhi N. The discovery of VLA-4 antagonists. Curr Top Med Chem 2004, 4(14):1509–1523. 23. Bennett JS. Novel platelet inhibitors. Annu Rev Biomed Eng 2001, 52:161–184. 24. Cai W, Chen X. Anti-angiogenic cancer therapy based on integrin alphavbeta3 antagonism. Anticancer Agents Med Chem 2006, 6(5):407–428. 25. Newby LK, Califf RM, White HD, Harrington RA, Van de Werf F, Granger CB, Simes RJ, Hasselblad V, Armstrong PW. The failure of orally administered glycoprotein IIb/IIIa inhibitors to prevent recurrent cardiac events. Am J Med 2002, 112(8):647–658. 26. Stupack DG, Cheresh DA. Integrins and angiogenesis. Curr Top Dev Biol 2004, 64:207–238. 27. Li ZB, Cai W, Cao Q, Chen K, Wu Z, He L, Chen X. (64)Cu-labeled tetrameric and octameric RGD peptides for small-animal PET of tumor alpha(v)beta(3) integrin expression. J Nucl Med 2007, 48(7):1162–1171. 28. Dijkgraaf I, Kruijtzer JA, Frielink C, Soede AC, Hilbers HW, Oyen WJ, Corstens FH, Liskamp RM, Boerman OC. Synthesis and biological evaluation of potent alphavbeta3-integrin receptor antagonists. Nucl Med Biol 2006, 33(8):953–961. 29. Burke PA, DeNardo SJ, Miers LA, Lamborn KR, Matzku S, DeNardo GL. Cilengitide targeting of alpha(v)beta(3) integrin receptor synergizes with radioimmunotherapy to increase efficacy and apoptosis in breast cancer xenografts. Cancer Res 2002, 62(15):4263–4272. 30. Tucker GC. Integrins: molecular targets in cancer therapy. Curr Oncol Rep 2006, 8(2):96–103. 31. Chen X. Multimodality imaging of tumor integrin alphavbeta3 expression. Mini Rev Med Chem 2006, 6(2):227–234. 32. Zhou Y, Drummond DC, Zou H, Hayes ME, Adams GP, Kirpotin DB, Marks JD. Impact of Single-chain Fv antibody fragment affinity on nanoparticle targeting of epidermal growth factor receptor-expressing tumor cells. J Mol Biol 2007, 371(4):934–947. 33. Kehoe JW, Kay BK. Filamentous phage display in the new millennium. Chem Rev 2005, 105(11):4056–4072. 34. Jostock T, Dubel S. Screening of molecular repertoires by microbial surface display. Comb Chem High Throughput Screen 2005, 8(2):127–133. 35. Huang BC, Liu R. Comparison of mRNA-displaybased selections using synthetic peptide and natural protein libraries. Biochemistry 2007, 46(35:):10102–10112.

332

36. Pasqualini R, Ruoslahti E. Organ targeting in vivo using phage display peptide libraries. Natura 1996, 380(6572):364–366. 37. Laakkonen P, Porkka K, Hoffman JA, Ruoslahti E. A tumor-homing peptide with a targeting specificity related to lymphatic vessels. Nat Med 2002, 8(7):751–755. 38. Arap W, Haedicke W, Bernasconi M, Kain R, Rajotte D, Krajewski S, Ellerby HM, Bredesen DE, Pasqualini R, Ruoslahti E. Targeting the prostate for destruction through a vascular address. Proc Natl Acad Sci U S A 2002, 99(3):1527–1531. 39. Lee SM, Lee EJ, Hong HY, Kwon MK, Kwon TH, Choi JY, Park RW, Kwon TG, Yoo ES, Yoon GS, Kim IS, Ruoslahti E, Lee BH. Targeting bladder tumor cells in vivo and in the urine with a peptide identified by phage display. Mol Cancer Res 2007, 5(1):11–19. 40. Pasqualini R, Koivunen E, Kain R, Lahdenranta J, Sakamoto M, Stryhn A, Ashmun RA, Shapiro LH, Arap W, Ruoslahti E. Aminopeptidase N is a receptor for tumor-homing peptides and a target for inhibiting angiogenesis. Cancer Res 2000, 60(3):722–727. 41. Essler M, Ruoslahti E. Molecular specialization of breast vasculature: a breast-homing phage-displayed peptide binds to aminopeptidase P in breast vasculature. Proc Natl Acad Sci U S A 2002, 99(4):2252–2257. 42. Pasqualini R, Arap W, McDonald DM. Probing the structural and molecular diversity of tumor vasculature. Trends Mol Med 2002, 8(12):563–571. 43. Roberts R, Makey DG, Seal US. Human transferrin. Molecular weight and sedimentation properties. J Biol Chem 1966, 241(21):4907–4913. 44. Carpenter G, Cohen S. Epidermal growth factor. J Biol Chem 1990, 265(14):7709–7712. 45. Holliger P, Hudson PJ. Engineered antibody fragments and the rise of single domains. Nat Biotechnol 2005, 23(9):1126–1136. 46. Karatan E, Merguerian M, Han Z, Scholle MD, Koide S, Kay BK. Molecular recognition properties of FN3 monobodies that bind the Src SH3 domain. Chem Biol 2004, 11(6):835–844. 47. Shen X, Valencia CA, Szostak JW, Dong B, Liu R. Scanning the human proteome for calmodulinbinding proteins. Proc Natl Acad Sci U S A 2005, 102(17):5969–5974. 48. Wadia JS, Dowdy SF. Protein transduction technology. Curr Opin Biotechnol 2002, 13(1):52–56. 49. Gupta B, Levchenko TS, Torchilin VP. Intracellular delivery of large molecules and small particles by cellpenetrating proteins and peptides. Adv Drug Deliv Rev 2005, 57(4):637–651. 50. Dietz GP, Bahr M. Delivery of bioactive molecules into the cell: the Trojan horse approach. Mol Cell Neurosci 2004, 27(2):85–131.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

51. Hawiger J. Noninvasive intracellular delivery of functional peptides and proteins. Curr Opin Chem Biol 1999, 3(1):89–94. ¨ 52. Tunnemann G, Martin RM, Haupt S, Patsch C, Edenhofer F, Cardoso MC. Cargo-dependent mode of uptake and bioavailability of TAT-containing proteins and peptides in living cells. FASEB J 2006, 20(11):1775–1784. 53. Bucci M, Gratton JP, Rudic RD, Acevedo L, Roviezzo F, Cirino G, Sessa WC. In vivo delivery of the caveolin-1 scaffolding domain inhibits nitric oxide synthesis and reduces inflammation. Nat Med 2000, 6(12):1362–1367. ¨ 54. May MJ, D’Acquisto F, Madge LA, Glockner J, Pober JS, Ghosh S. Selective inhibition of NF-kappaB activation by a peptide that blocks the interaction of NEMO with the IkappaB kinase complex. Science 2000, 289(5484):1550–1554. 55. Nagahara H, Vocero-Akbani AM, Snyder EL, Ho A, Latham DG, Lissy NA, Becker-Hapak M, Ezhevsky SA, Dowdy SF. Transduction of full-length TAT fusion proteins into mammalian cells: TATp27Kip1 induces cell migration. Nat Med 1998, 4(12):1449–1452. 56. Schwarze SR, Ho A, Vocero-Akbani A, Dowdy SF. In vivo protein transduction: delivery of a biologically active protein into the mouse. Science 1999, 285(5433):1569–1572. 57. Jo D, Nashabi A, Doxsee C, Lin Q, Unutmaz D, Chen J, Ruley HE. Epigenetic regulation of gene structure and function with a cell-permeable Cre recombinase. Nat Biotechnol 2001, 19(10):929–933. 58. Trehin R, Merkle HP. Chances and pitfalls of cell penetrating peptides for cellular drug delivery. Eur J Pharm Biopharm 2004, 58(2):209–223. 59. Leifert JA, Harkins S, Whitton JL. Full-length proteins attached to the HIV tat protein transduction domain are neither transduced between cells, nor exhibit enhanced immunogenicity. Gene Ther 2002, 9(21):1422–1428. 60. Juliano RL. Peptide-oligonucleotide conjugates for the delivery of antisense and siRNA. Curr Opin Mol Ther 2005, 7(2):132–136. 61. Fisher AA, et al. Evaluating the specificity of antisense oligonucleotide conjugates. A DNA array analysis. J Biol Chem 2002, 277(25):22980–22984. 62. Astriab-Fisher A, et al. Antisense inhibition of Pglycoprotein expression using peptide-oligonucleotide conjugates. Biochem Pharmacol 2000, 60(1):83–90. ¨ 63. Pooga M, Soomets U, Hallbrink M, Valkna A, Saar K, Rezaei K, Kahl U, Hao JX, Xu XJ, Wiesenfeld-Hallin ¨ Z, Hokfelt T, Bartfai T, Langel U. Cell penetrating PNA constructs regulate galanin receptor levels and modify pain transmission in vivo. Nat Biotechnol 1998, 16(9):857–861.

Vo lu me 1, May /Ju n e 2009

64. Moulton HM, Hase MC, Smith KM, Iversen PL. HIV Tat peptide enhances cellular delivery of antisense morpholino oligomers. Antisense Nucleic Acid Drug Dev 2003, 13(1):31–43. 65. Abes S, Moulton H, Turner J, Clair P, Richard JP, Iversen P, Gait MJ, Lebleu B. Peptide-based delivery of nucleic acids: design, mechanism of uptake and applications to splice-correcting oligonucleotides. Biochem Soc Trans 2007, 35(Pt 1):53–55. 66. Torchilin VP, Rammohan R, Weissig V, Levchenko TS. TAT peptide on the surface of liposomes affords their efficient intracellular delivery even at low temperature and in the presence of metabolic inhibitors. Proc Natl Acad Sci U S A 2001, 98(15):8786–8791. 67. Zhao M, Kircher MF, Josephson L, Weissleder R. Differential conjugation of tat peptide to superparamagnetic nanoparticles and its effect on cellular uptake. Bioconjug Chem 2002, 13(4):840–844. 68. Perret E, Lakkaraju A, Deborde S, Schreiner R, Rodriguez-Boulan E. Evolving endosomes: how many varieties and why? Curr Opin Cell Biol 2005, 17(4):423–434. 69. Sandvig K, van Deurs B. Delivery into cells: lessons learned from plant and bacterial toxins. Gene Ther 2005, 12(11):865–872. 70. Sieczkarski SB, Whittaker GR. Viral entry. Curr Top Microbiol Immunol 2005, 285:1–23. 71. Marsh M, Helenius A. Virus entry: open sesame. Cell 2006, 124(4):729–740. 72. Derossi D, Calvet S, Trembleau A, Brunissen A, Chassaing G, Prochiantz A. Cell internalization of the third helix of the Antennapedia homeodomain is receptor-independent. J Biol Chem 1996, 271(30:):18188–18193. 73. Richard JP, Melikov K, Vives E, Ramos C, Verbeure B, Gait MJ, Chernomordik LV, Lebleu B. Cell-penetrating peptides. A reevaluation of the mechanism of cellular uptake. J Biol Chem 2003, 278(1):585–590. 74. Tyagi M, Rusnati M, Presta M, Giacca M. Internalization of HIV-1 tat requires cell surface heparan sulfate proteoglycans. J Biol Chem 2001, 276(5):3254–3261. ¨ 75. Fischer R, Kohler K, Fotin-Mleczek M, Brock R. A stepwise dissection of the intracellular fate of cationic cell-penetrating peptides. J Biol Chem 2004, 279(13):12625–12635. 76. Wadia JS, Stan RV, Dowdy SF. Transducible TATHA fusogenic peptide enhances escape of TAT-fusion proteins after lipid raft macropinocytosis. Nat Med 2004, 10(3):310–315. 77. Silhol M, Tyagi M, Giacca M, Lebleu B, Viv`es E. Different mechanisms for cellular internalization of the HIV-1 Tat-derived cell penetrating peptide and recombinant proteins fused to Tat. Eur J Biochem 2002, 269(2):494–501.

 2009 Jo h n Wiley & So n s, In c.

333

Advanced Review

www.wiley.com/wires/nanomed

78. Christiaens B, Symoens S, Verheyden S, Engelborghs Y, Joliot A, Prochiantz A, Vandekerckhove J, Rosseneu M, Vanloo B. Tryptophan fluorescence study of the interaction of penetratin peptides with model membranes. Eur J Biochem 2002, 269(12):2918–2926. 79. Thor´en PE, Persson D, Lincoln P, Nord´en B. Membrane destabilizing properties of cell-penetrating peptides. Biophys Chem 2005, 114(2–3):169–179. 80. Lee HJ, Pardridge WM. Pharmacokinetics and delivery of tat and tat-protein conjugates to tissues in vivo. Bioconjug Chem 2001, 12(6):995–999. 81. Jain M, Chauhan SC, Singh AP, Venkatraman G, Colcher D, Batra SK. Penetratin improves tumor retention of single-chain antibodies: a novel step toward optimization of radioimmunotherapy of solid tumors. Cancer Res 2005, 65(17):7840–7846. ¨ 82. Niesner U, Halin C, Lozzi L, Gunthert M, Neri P, Wunderli-Allenspach H, Zardi L, Neri D. Quantitation of the tumor-targeting properties of antibody fragments conjugated to cell-permeating HIV-1 TAT peptides. Bioconjug Chem 2002, 13(4):729–736. 83. Amantana A, Moulton HM, Cate ML, Reddy MT, Whitehead T, Hassinger JN, Youngblood DS, Iversen PL. Pharmacokinetics, biodistribution, stability and toxicity of a cell-penetrating peptidemorpholino oligomer conjugate. Bioconjug Chem 2007, 18(4):1325–1331. 84. Temming K, Lacombe M, van der Hoeven P, Prakash J, Gonzalo T, Dijkers EC, Orfi L, K´eri G, Poelstra K, Molema G, Kok RJ. Delivery of the p38 MAPkinase inhibitor SB202190 to angiogenic endothelial cells: development of novel RGD-equipped and PEGylated drug-albumin conjugates using platinum(II)based drug linker technology. Bioconjug Chem 2006, 17(5):1246–1255. 85. Temming K, Meyer DL, Zabinski R, Senter PD, Poelstra K, Molema G, Kok RJ. (Molecular Pharmaceutics, not Molecular Pharmacology) Improved efficacy of alphavbeta3-targeted albumin conjugates by conjugation of a novel auristatin derivative. Mol Pharmacol 2007, 4(5):686–694. 86. Hariton-Gazal E, Rosenbluh J, Zakai N, Fridkin G, Brack-Werner R, Wolff H, Devaux C, Gilon C, Loyter A. Functional analysis of backbone cyclic peptides bearing the arm domain of the HIV-1 rev protein: characterization of the karyophilic properties and inhibition of rev-induced gene expression. Biochemistry 2005, 44(34:):11555–11566. 87. Morishita M, et al. A novel approach using functional peptides for efficient intestinal absorption of insulin. J Control Release 2007, 118(2):177–184. 88. Zhang C, et al. Specific targeting of tumor angiogenesis by RGD-conjugated ultrasmall superparamagnetic iron oxide particles using a clinical 1.5-T magnetic resonance scanner. Cancer Res 2007, 67(4):1555–1562.

334

89. Schmieder AH, et al. Molecular MR imaging of melanoma angiogenesis with alphanubeta3-targeted paramagnetic nanoparticles. Magn Reson Med 2005, 53(3):621–627. 90. Lewin M, Carlesso N, Tung CH, Tang XW, Cory D, Scadden DT, Weissleder R. Tat peptide-derivatized magnetic nanoparticles allow in vivo tracking and recovery of progenitor cells. Nat Biotechnol 2000, 18(4):410–414. 91. Pittet MJ, Swirski FK, Reynolds F, Josephson L, Weissleder R. Labeling of immune cells for in vivo imaging using magnetofluorescent nanoparticles. Nat Protoc 2006, 1(1):73–79. 92. Sahoo SK, Ma W, Labhasetwar V. Efficacy of transferrin-conjugated paclitaxel-loaded nanoparticles in a murine model of prostate cancer. Int J Cancer 2004, 112(2):335–340. 93. Heidel JD, Yu Z, Liu JY, Rele SM, Liang Y, Zeidan RK, Kornbrust DJ, Davis ME. Administration in non-human primates of escalating intravenous doses of targeted nanoparticles containing ribonucleotide reductase subunit M2 siRNA. Proc Natl Acad Sci U S A 2007, 104(14):5715–5721. 94. Kompella UB, Sundaram S, Raghava S, Escobar ER. Luteinizing hormone-releasing hormone agonist and transferrin functionalizations enhance nanoparticle delivery in a novel bovine ex vivo eye model. Mol Vis 2006, 12:1185–1198. 95. Yong KT, Qian J, Roy I, Lee HH, Bergey EJ, Tramposch KM, He S, Swihart MT, Maitra A, Prasad PN. Quantum rod bioconjugates as targeted probes for confocal and two-photon fluorescence imaging of cancer cells. Nano Lett 2007, 7(3):761–765. 96. Kim WJ, Yockman JW, Jeong JH, Christensen LV, Lee M, Kim YH, Kim SW. Anti-angiogenic inhibition of tumor growth by systemic delivery of PEI-g-PEGRGD/pCMV-sFlt-1 complexes in tumor-bearing mice. J Control Release 2006, 114(3):381–388. 97. Ding H, Prodinger WM, Kopecek J. Identification of CD21-binding peptides with phage display and investigation of binding properties of HPMA copolymer-peptide conjugates. Bioconjug Chem 2006, 17(2):514–523. 98. Mitra A, Coleman T, Borgman M, Nan A, Ghandehari H, Line BR. Polymeric conjugates of monoand bi-cyclic alphaVbeta3 binding peptides for tumor targeting. J Control Release 2006, 114(2):175–183. ¨ 99. Kunath K, Merdan T, Hegener O, Haberlein H, Kissel T. Integrin targeting using RGD-PEI conjugates for in vitro gene transfer. J Gene Med 2003, 5(7):588–599. 100. Kursa M, Walker GF, Roessler V, Ogris M, Roedl W, Kircheis R, Wagner E. Novel shielded transferrinpolyethylene glycol-polyethylenimine/DNA complexes for systemic tumor-targeted gene transfer. Bioconjug Chem 2003, 14(1):222–231.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Cell targetting and CPPs

101. Kleemann E, Neu M, Jekel N, Fink L, Schmehl T, Gessler T, Seeger W, Kissel T. Nano-carriers for DNA delivery to the lung based upon a TAT-derived peptide covalently coupled to PEG-PEI. J Control Release 2005, 109(1–3):299–316.

114. Kale AA, Torchilin VP. Enhanced transfection of tumor cells in vivo using ”Smart” pH-sensitive TATmodified pegylated liposomes. J Drug Target 2007, 15(7):538–545.

102. Moffatt S, Wiehle S, Cristiano RJ. Tumor-specific gene delivery mediated by a novel peptidepolyethylenimine-DNA polyplex targeting aminopeptidase N/CD13. Hum Gene Ther 2005, 16(1):57–67.

115. Chen X, Hou Y, Tohme M, Park R, Khankaldyyan V, Gonzales-Gomez I, Bading JR, Laug WE, Conti PS. Pegylated Arg-Gly-Asp peptide: 64Cu labeling and PET imaging of brain tumor alphavbeta3-integrin expression. J Nucl Med 2004, 45(10):1776–1783.

103. Hildebrandt IJ, Iyer M, Wagner E, Gambhir SS. Optical imaging of transferrin targeted PEI/DNA complexes in living subjects. Gene Ther 2003, 10(9):758–764.

116. Kim JW, Lee HS. Tumor targeting by doxorubicinRGD-4C peptide conjugate in an orthotopic mouse hepatoma model. Int J Mol Med 2004, 14(4):529–535.

104. Nori A, Jensen KD, Tijerina M, Kopeckova´ P, Kopecek J. Subcellular trafficking of HPMA copolymer-tat conjugates in human ovarian carcinoma cells. J Control Release 2003, 91(1–2):53–59.

117. Janssen ML, Oyen WJ, Dijkgraaf I, Massuger LF, Frielink C, Edwards DS, Rajopadhye M, Boonstra H, Corstens FH, Boerman OC. Tumor targeting with radiolabeled α(v) beta(3) integrin binding peptides in a nude mouse model. Cancer Res 2002, 62(21):6146–6151.

105. Segura T, Hubbell JA. Synthesis and in vitro characterization of an ABC triblock copolymer for siRNA delivery. Bioconjug Chem 2007, 18(3):736–745. 106. Sethuraman VA, Bae YH. TAT peptide-based micelle system for potential active targeting of anti-cancer agents to acidic solid tumors. J Control Release 2007, 118(2):216–224. 107. Xiong XB, Huang Y, Lu WL, Zhang X, Zhang H, Nagai T, Zhang Q. Intracellular delivery of doxorubicin with RGD-modified sterically stabilized liposomes for an improved antitumor efficacy: in vitro and in vivo. J Pharm Sci 2005, 94(8):1782–1793. 108. Koning GA, Schiffelers RM, Wauben MH, Kok RJ, Mastrobattista E, Molema G, ten Hagen TL, Storm G. Targeting of angiogenic endothelial cells at sites of inflammation by dexamethasone phosphate-containing RGD peptide liposomes inhibits experimental arthritis. Arthritis Rheum 2006, 54(4):1198–1208. 109. Koning GA, Fretz MM, Woroniecka U, Storm G, Krijger GC. Targeting liposomes to tumor endothelial cells for neutron capture therapy. Appl Radiat Isot 2004, 61(5):963–967. 110. Dubey PK, Mishra V, Jain S, Mahor S, Vyas SP. Liposomes modified with cyclic RGD peptide for tumor targeting. J Drug Target 2004, 12(5):257–264. 111. Suresh Reddy J, Venkateswarlu V, Koning GA. Radioprotective effect of transferrin targeted citicoline liposomes. J Drug Target 2006, 14(1):13–19. 112. Moreira JN, Hansen CB, Gaspar R, Allen TM. A growth factor antagonist as a targeting agent for sterically stabilized liposomes in human small cell lung cancer. Biochim Biophys Acta 2001, 1514(2):303–317. ¨ 113. Kullberg EB, Wei Q, Capala J, Giusti V, Malmstrom PU, Gedda L. EGF-receptor targeted liposomes with boronated acridine: growth inhibition of cultured glioma cells after neutron irradiation. Int J Radiat Biol 2005, 81(8):621–629.

Vo lu me 1, May /Ju n e 2009

118. Chen X, Plasencia C, Hou Y, Neamati N. Synthesis and biological evaluation of dimeric RGD peptidepaclitaxel conjugate as a model for integrin-targeted drug delivery. J Med Chem 2005, 48(4):1098–1106. 119. Wu Y, Zhang X, Xiong Z, Cheng Z, Fisher DR, Liu S, Gambhir SS, Chen X. microPET imaging of glioma integrin alphavbeta3 expression using (64)Culabeled tetrameric RGD peptide. J Nucl Med 2005, 46(10):1707–1718. 120. Ye Y, Bloch S, Xu B, Achilefu S. Design, synthesis, and evaluation of near infrared fluorescent multimeric RGD peptides for targeting tumors. J Med Chem 2006, 49(7):2268–2275. 121. Liu S, Hsieh WY, Jiang Y, Kim YS, Sreerama SG, Chen X, Jia B, Wang F. Evaluation of a (99m)Tc-labeled cyclic RGD tetramer for noninvasive imaging integrin alpha(v)beta3-positive breast cancer. Bioconjug Chem 2007, 18(2):438–446. 122. Alves S, Correia JD, Gano L, Rold TL, Prasanphanich A, Haubner R, Rupprich M, Alberto R, Decristoforo C, Santos I, Smith CJ. In vitro and in vivo evaluation of a novel 99mTc(CO)3-pyrazolyl conjugate of cyclo-(Arg-Gly-Asp-d-Tyr-Lys). Bioconjug Chem 2007, 18(2):530–537. ¨ 123. Garc´ıa-Garayoa E, Maes V, Blauenstein P, Blanc A, Hohn A, Tourw´e D, Schubiger PA. Double-stabilized neurotensin analogues as potential radiopharmaceuticals for NTR-positive tumors. Nucl Med Biol 2006, 33(4):495–503. 124. McQuade P, Knight LC, Welch MJ. Evaluation of 64Cu- and 125I-radiolabeled bitistatin as potential agents for targeting alpha v beta 3 integrins in tumor angiogenesis. Bioconjug Chem 2004, 15(5):988–996. 125. Lim CJ, Shen WC. Transferrin-oligomers as potential carriers in anticancer drug delivery. Pharm Res 2004, 21(11):1985–1992.

 2009 Jo h n Wiley & So n s, In c.

335

Advanced Review

Nanotechnology for bone materials Nhiem Tran1 and Thomas J. Webster2∗ It has been established that for orthopedic-related research, nanomaterials (materials defined as those with constituent dimensions less than 100 nm in at least one direction) have superior properties compared to conventional counterparts. This review summarizes studies that have demonstrated enhanced in vitro and in vivo osteoblast (bone-forming cells) functions (such as adhesion, proliferation, synthesis of bone-related proteins, and deposition of calcium-containing mineral) on nanostructured metals, ceramics, polymers, and composites thereof compared to currently used implants. These results strongly imply that nanomaterials may improve osseointegration, which is crucial for long-term implant efficacy. This review also focuses on novel drug-carrying magnetic nanoparticles designed to treat various bone diseases (such as osteoporosis). Although further investigation of the in vivo responses and toxicity of these novel nanomaterials pertinent for orthopedic applications are needed, nanotechnology clearly has already demonstrated the ability to produce better bone implants and therefore should be further investigated.  2009 John Wiley & Sons, Inc. WIREs Nanomed Nanobiotechnol 2009 1 336–351

NANOTECHNOLOGY: A NEW APPROACH TO IMPROVE ORTHOPEDIC IMPLANTS

B

one is the main component of the human skeletal system, and as such serves a vital function in our daily lives. Bones exist in various shapes and have complex internal and external structures which contribute to their light weight and high strength. In the body, bones serve the following three functions:1 (1) provide mechanical support, as it is the site of muscle attachment for movement; (2) protect various organs, including bone marrow filled with nutrients; and (3) provide a metabolic function (storing calcium, phosphorus, and other essential ions to be used by the body when needed). Depending on the anatomical location, the properties and structure of the bone can significantly differ. However, there are some commonalities between all bone types. Specifically, bone is a natural nanostructured (that is, a material with constituent features less than 100 nm in at least one dimension) composite composed of

∗

Correspondence to: thomas [email protected]

1 Department

of Physics, Division of Engineering and Department of Orthopedics, Brown University, Providence, RI 02917, USA 2 Division

of Engineering and Department of Orthopedics, Brown University, Providence, RI 02917, USA DOI: 10.1002/wnan.023

336

organic compounds (mainly collagen) reinforced with inorganic ones (hydroxyapatite). It is this natural nanostructure that nanotechnology aims to emulate for orthopedic applications. However, as with any organ, bone diseases are very common in our society. In fact, annually, an estimated 1.5 million individuals in the United States suffer from bone fractures.2 Among the diseases leading to bone fractures, osteoporosis is the leading cause in both males and females of all ages. Osteoporosis is a silent disease which is characterized by reduced bone mineral density and a deteriorated structure of the bone tissue.3 A common treatment for bone fractures is the implantation of a mostly metallic orthopedic prosthetic. These prosthetics help heal bone nonunions and allow patients to partially regain function. However, today’s orthopedic implant materials do not allow patients to return to their normal, daily active lifestyles that they had before fracture. Specifically, it has been reported that the average lifetime of orthopedic implants is only 10–15 years.4 This means that those who are young and receive a traditional orthopedic implant will have to undergo several more painful and expensive surgeries to replace such a failed orthopedic implant. Although controversial, several factors lead to implant failures such as incomplete,

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

Digital instruments NanoScope Scan size 5.000 µm Scan rate 1.001 hz Number of samples 256 Image Data Height Data scale 250.0 m Numerous Nanostructured Surface Features view angle light angle

FIGURE 1 | Representative AFM image of cortical

bovine bone. Numerous nanostructured features on the surface of cortical bovine bone are visible. Root-mean-square values from AFM for 5 µm × 5 µm and 25 µm × 25 µm scans were 32 and 25 nm, respectively. (Reprinted with permission from Ref 6. Copyright 2005 Institute of Physics Publishing (IOP)).

1 2 3 4 µm

prolonged osseointegration5–7 (i.e., lack of bonding of an orthopedic implant to juxtaposed bone) and/or severe stress shielding (due to differences in mechanical properties between an implant and the surrounding bone8,9 ). Therefore, a great amount of effort from biomedical researchers worldwide have focused on improving the design and manufacture of orthopedic implants to last longer in the body. While many have attempted to alter orthopedic implant chemistry (from metals to ceramics to polymers), recent discoveries have highlighted that nanotechnology may universally improve all materials used for regrowing bone. This approach has been used on a wide range of materials (such as metals, ceramics, polymers, and composites), in which either nanostructured surface features or constituent nanomaterials (including grains, fibers, or particles with at least one dimension from 1 to 100 nm) have been implemented. These nanomaterials have shown superior properties compared to their conventional (or micron structured) counterparts due to their distinctive nanoscale features and novel physical properties that ensue.10,11 As mentioned, the novelty of nanotechnology is that it creates materials that mimic the natural nanostructure of our tissues (Figure1). Specifically, natural bone is composed of three levels of hierarchy:12 (1) the nanostructure (a few nanometers to a few hundred nanometers), such as noncollagenous organic proteins, fibrillar collagen, embedded Vo lu me 1, May /Ju n e 2009

x 1,000 µm/div z 250,000 m/div

calcium phosphate crystals, etc.; (2) the microstructure (1–500µm), such as lamellae, osteons, and Haversian systems; and (3) the macrostructure, such as cancellous and cortical bone. These structures assemble into unique heterogeneous and anisotropic bone that has yet to be accurately duplicated in today’s bone implants. Of relevance to the merging of nanotechnology and orthopedics is the realization that bone is a natural nanostructured composite material composed of intertwined inorganic (bone apatite) and organic compounds (mainly collagen).13 Although this has been known for quite some time, only today has there been an emphasis on duplicating the same nanofeatures in bone that osteoblasts (boneforming cells) interact with and synthesize themselves. Specifically, type I collagen contributes approximately 90% of the organic phase of bone and noncollagenous proteins and ground substances the rest (10%).152 As for the inorganic phase of bone, various forms of calcium phosphates (most notably, crystalline hydroxyapatite, Ca10 (PO4 )6 (OH)2 or HA) are the primary mineral components. Since bone is a hierarchical material with the lowest level belonging to the nanoscale range, materials with nanometer structures (as only nanotechnology can create) appear as natural choices for creating better bone implants. As will be described, metallic, ceramic, polymeric, and composites thereof have all been tested for their ability to serve as novel orthopedic materials.

 2009 Jo h n Wiley & So n s, In c.

337

Advanced Review

www.wiley.com/wires/nanomed

Nanophase Ti

Nanophase CoCrMo

Nanophase Ti6Al4V

Conventional Ti

Conventional CoCrMo

Conventional Ti6Al4V

NANOMATERIALS IN BONE-RELATED RESEARCH Metals Metals have been widely used as orthopedic implant materials mainly because of their superior mechanical properties which allow for load-bearing situations.153 They are used as bone replacement materials in hip prostheses, dental implants, and so on. The most commonly used metals in orthopedics are stainless steel,14 cobalt chrome alloys,15 titanium, and titanium alloys.15,16 Despite the advantage of mechanical properties of metals, metals often fail (i.e., permanently separate from bone) after 10–15 years of implantation into the human body. Researchers explain this failure to be the consequence of incomplete, prolonged osseointegration between the implant and surrounding bone.5–9 Other disadvantages of metallic orthopedic materials include the need for further operation to remove temporary implants such as plates, pins, and screws, and negative tissue responses to the ions released from metallic implant materials.17,18,154 However, these disadvantages common for conventional (or micron-structured) metals may not be the same for nanostructured metals. Specifically, several studies have reported increased adhesion and functions of osteoblasts on metallic nanoscale surface features (compared to what is currently implanted),10,11,19,20 thus giving a strong potential improvement for the prolonged osseointegration of nanostructured metals to the surrounding bone. For example, research groups have reported increased osteoblast functions (such as adhesion, proliferation, and deposition of calcium-containing mineral) on the following nanophase compared to 338

FIGURE 2 | Scanning electron microscopy images of Ti, CoCrMo, and Ti6A14V compacts. Increased nanostructured surface roughness was observed on nanophase compared to conventional Ti, CoCrMo, and Ti6A14V. Bar = 1 µm for nanophase compacts and 10 µm for conventional ones. (Reprinted with permission from Ref 21. Copyright 2004 Elsevier).

conventional metals: Ti, Ti6Al4V, and CoCrMo20 (Figure 2). Nanometer- and micron-sized powders of Ti, Ti6 Al4 V, and CoCrMo have been separately pressed into model implant surfaces, and osteoblast adhesion has been observed after 1 and 3 h. The results showed that osteoblast adhesion was significantly greater on nanophase Ti, Ti6Al4V, and CoCrMo when compared to conventional micron-grain-size Ti, Ti6Al4V, and CoCrMo. Interestingly, it was also observed that osteoblast adhesion occurred primarily at metal particle boundaries. Thus, these results indicated that increased osteoblast adhesion might be due to the greater number of defect boundaries at the surface of nanophase materials, creating unique surface energy attracting the adsorption of certain proteins important for osteoblast adhesion. Since the adhesion of osteoblasts is a prerequisite for subsequent functions (such as the deposition of calcium-containing minerals), this study implied further improved functions of osteoblasts on nanophase Ti, Ti6Al4V, and CoCrMo (which was later confirmed).6 However, there are other ways to create nanostructured metals besides the aforementioned compacting of metallic nanostructured particles. For example, a recent study by Yao et al. demonstrated another way to elevate the performance of metallic orthopedic materials by using an anodization process.21 In that study, the surface of Ti was anodized to create nanotube-like pores, which possessed higher surface energy and wettability compared to unanodized Ti. This nanoscaled Ti greatly improved osteoblast adhesion.21 Following this research direction, Sirivisoot et al. grew carbon nanotubes (CNTs) on the surface of the nanotubular anodized Ti in

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

an effort to create in situ orthopedic biosensors to detect and control bone growth22 (Figure 3). As an additional advantage, osteoblast long-term function studies demonstrated significantly higher alkaline phosphate activity and calcium deposition by osteoblasts on these novel orthopedic sensors composed of CNTs grown out of anodized nanotubular Ti compared to conventional Ti. Several other recent studies have indicated an improvement of bone cell functions on nanophase materials.10,11 However, some of these results have been ambiguous concerning the contribution of nanometer surface roughness since those nanophase materials simultaneously changed in roughness and surface chemistry.23–26 Since nanotechnology is defined as novel material properties dependent on the nanometer size scale, it is important to separate chemical from nanostructured roughness effects. Khang et al. provided insight into the role of nanometer surface features alone (with no change in surface chemistry) in enhancing tissue growth by comparing (a)

bone cell functions on three types of materials differing only in surface structures (i.e., flat, nanometer, and submicron).27 Briefly, nanometer and submicron surface patterns were produced on currently implanted Ti via e-beam deposition of Ti. Creating patterns of nanostructured features on metallic orthopedic implant surfaces has been a focus area of numerous researchers to emulate the anisotropic alignment of collagen and hydroxyapatite in long bones of the body (such as the femur). The analysis of osteoblast adhesion on such highly aligned patterns of flat/submicron and flat/nanometer patterned Ti surfaces showed that cell densities were significantly higher on both the nanometer and submicron surfaces compared to flat Ti surfaces. This study further demonstrated that nanometer roughness had the highest efficiency for increasing both surface energy and osteoblast adhesion compared to the altered pattern width of surface features, and therefore suggested an important role of nanometer and submicron surface features on the future design of improved Ti-based implants.

(b)

1 µm

(c)

1 µm

(d)

200 nm

(e)

200 nm

(f)

1 µm

Vo lu me 1, May /Ju n e 2009

100 nm

 2009 Jo h n Wiley & So n s, In c.

FIGURE 3 | SEM micrographs of:

(a) unanodized Ti, (b) anodized Ti without CNTs, (c) lower and (d) higher magnification of CNTs grown from the nanotubes of anodized Ti without a Co catalyst, and (e) lower and (f) higher magnification of CNTs grown from the nanotubes of anodized Ti surface with a Co catalyst. (Reprinted with permission from Ref 22. Copyright 2005 Institute of Physics Publishing (IOP)).

339

Advanced Review

www.wiley.com/wires/nanomed

Ceramics However, it is not just nanostructured metals that promote bone cell functions. In fact, one of the earliest studies on the impact of nanotechnology for orthopedic applications investigated osteoblast function on nanostructured ceramics and found that nanostructured ceramics promoted bone cell functions compared to traditional, nano-smooth ceramics.10,11 Ceramics have been used extensively in orthopedics owing to their well-known biocompatibility properties with bone cells and tissues.28–31 Metallic oxides (e.g., alumina, zirconia, titania), calcium phosphates (e.g., hydroxyapatite (HA), tricalcium phosphate (TCP), calcium tetraphosphate (Ca4 P2 O9 )),32 and glass ceramics (e.g., Bioglass and Ceravital)33 are among those commonly used in orthopedic tissue engineering applications. These ceramics are considered bioactive because of their surface properties that support bone cell adhesion, proliferation, and differentiation. Specific ceramics (such as HA and TCP) have similar chemistry to the mineral phase of natural bone. Consequently, their reaction with physiological fluids creates strong bonds to hard and soft tissues, thereby increasing osseointegration between implants and bone. Moreover, these ceramics are degradable and their dissolution rate depends on crystallinity. Therefore, the degradation of many ceramics can be controlled to match the rate of new bone growth. It is well known that the degradation time of amorphous calcium phosphate is much shorter than that of crystalline HA.34 In addition, another level of control can be gained by the understanding that nanocrystalline calcium phosphates degrade faster than micron (or conventional) grain size calcium phosphates.34 It has long been known that conventional ceramics (such as alumina, titania, and HA with grain sizes greater than 100 nm) possess good biocompatibility properties.28–31 However, similar to metals, clinical applications of these ceramics still encounter many difficulties because of insufficient prolonged bonding to the juxtaposed bone. In addition, applications for large bone defects are not feasible because of the natural brittleness of ceramics. Therefore, developing novel ceramic materials that can promote and sustain osseointegration with the surrounding bone is necessary for improved orthopedic implant applications. The first in vitro study highlighting the influence of ceramic nanometer grain size on bone cell adhesion was reported in 1999.10 Compared to micronsized conventional substrates, osteoblast adhesion was significantly higher on nanophase alumina and titania substrates. Since an increase in osteoblast 340

adhesion was observed for both nanophase alumina and titania, this study implied that the enhanced osteoblast adhesion was independent of surface chemistry and was dependent only on the surface topography and subsequent roughness of ceramics; this is the same versatile trend previously described for nanostructured compared to conventional metals. Further evidence of enhanced functions of osteoblasts (such as proliferation, alkaline phosphatase activity, and calcium deposition) on nanophase ceramics has also been provided.11 Osteoblast proliferation results after 5 days showed higher cell densities on nanophase titania, alumina, and HA compared to the respective conventional grain size ceramics. Moreover, calcium content in the extracellular matrix of osteoblasts cultured on nanophase alumina, titania, and HA was 4, 6, and 2 times greater than on the respective conventional ceramics. Overall, the results of these studies for the first time indicated that nanophase ceramics, with their unique surface properties, can enhance bonding of orthopedic/dental implants to juxtaposed bone and therefore possibly improve bone implant efficacy. The mechanisms underlying the superior bone growing properties of nanophase ceramics to regenerate bone compared to the respective conventional formulations have been reported to involve vitronectin (a key protein that mediates osteoblast adhesion).35 It is known that cell adhesion, an important prerequisite for further cell functions, involves initial protein adsorption since osteoblast adhesion is greatly reduced in the absence of serum-containing proteins regardless of ceramic grain size.36 Owing to the increased wettability of nanostructured ceramics, results showed that vitronectin (a hydrophilic protein) adsorbed in the highest concentration on nanophase alumina compared to conventional alumina. Decreased adsorption of apolipoprotein A–I and/or increased adsorption of calcium on nanophase alumina were additional possible causes of the observed enhanced adsorption of vitronectin on nanophase alumina.35 Moreover, that study provided evidence of different conformations of vitronectin adsorbed on nanophase compared to conventional alumina, resulting in an increased unfolding of the cell-adhesive macromolecule to expose epitopes (such as arginine-glycine-aspartic acid or RGD) recognized by specific cell-membrane receptors.

Polymers Among all of the materials mentioned for bone tissue engineering applications, polymers are often chosen (at least in part) because they have physical properties that closely resemble those of proteins in soft and

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

hard tissues. Polymers are easily manufactured into programmed shapes and structures. In addition, polymers can be modified chemically or functionalized via chemical and biochemical reactions.11 Synthetic polymers have attracted growing attention in tissue engineering applications37 since, in general, synthetic polymers can be fabricated and tailored to give a wide range of properties and often are free from concerns of immunogenicity.38 The ideal polymeric material should match the demands of particular orthopedic applications, and therefore should exhibit the following properties:38 (1) not evoke an inflammatory/toxic response disproportionate to its beneficial effect, (2) degrade after fulfilling its purpose and leave no harmful components, (3) be easily processed into the final product form, (4) have an acceptable shelf life, and (5) be easily sterilized. Controllable degradation time and mechanical properties that match that of the organic component of bone are crucial since implants and scaffolds need to have sufficient strength until the surrounding tissue has regenerated. Compared to other implant materials (such as metals and ceramics), polymers have some undeniable advantages. Their mechanical properties and degradation times can be more easily tailored than with other materials. They can also be injectable and hardened in situ. In addition, the wide range of polymer chemical properties offers a diverse number of functionalizable materials to interact with different cell types. However, the major disadvantage of using polymers is that they can never by themselves approximate the mechanical properties of bone. The most popular natural polymer for tissue engineering is collagen. Collagen is a fibrous, nanoscale protein and a major component of natural extracellular matrices.155 Owing to its attractive biological properties (such as biocompatibility), collagen has been used for scaffold fabrication.39–41 However, there are still several concerns over the use of collagen for orthopedic applications because of poor handling and poor mechanical properties, and because of this synthetic nanostructured polymeric mimics have been created.42 For scaffold applications, collagen has to be manufactured into a threedimensional porous structure. This porous structure provides critical functions for the scaffold,43 such as allowing migration of cultured cells into the scaffold, providing a very large surface area for the cells to interact with the scaffold, and also allowing nutrients to diffuse into the scaffold. Most of the synthetic degradable polymers studied for orthopedic applications belong to the poly(αhydroxy acid) family, including poly(lactic acid) (PLA) Vo lu me 1, May /Ju n e 2009

(also known as polylactide), poly(glycolic acid) (PGA) (also known as polyglycolide), and their copolymers such as poly(lactic-co-glycolic acid) (PLGA) (also known as polylactide-co-glycolide).44–46 These polymers are also some of the few polymers that have been approved by the Food and Drug Administration (FDA) for human use in various medical devices. The degradation rate of these polymers can be manipulated via changes in the ratio of polylactic acid to PGA, molecular weight, crystallinity, hydrophylicity, pH of the surrounding environment, as well as specimen size, geometry, porosity, surface properties, and the sterilization process.156 For example, under the same conditions, hydrophilic PGA degrades much faster than hydrophobic PLA in vivo or in aqueous solutions. Therefore, PGA is frequently used as a component to fabricate a copolymer with PLA to control the degradation rate of the copolymer. Just like the ceramics and metals mentioned above, studies have highlighted that the efficacy of polymeric materials for bone regeneration can be improved through the use of nanotechnology. Recent research results have indeed demonstrated that the adsorption and conformation of proteins (such as fibronectin and vitronectin), which regulate osteoblast adhesion and other functions, are enhanced on nanophase compared to conventional polymers.6,35,36 For example, Wei and Ma reported the selective adsorption of proteins, including fibronectin and vitronectin, on 3D nanofibrous polymer scaffolds.47 In that study, nanoparticulate hydroxyapatite/poly(Llactic acid) (NHAP/PLLA) composite scaffolds were prepared using phase separation techniques and protein adsorption compared to micron particulate hydroxyapatite/poly(L-lactic acid) (MHAP/PLLA) scaffolds. The results showed significantly higher protein adsorption on NHAP/PLLA scaffolds of higher HA content. Consequently, this led to enhanced osteoblast functions (such as mineral deposition) on nanofibrous polymer scaffolds. In another study, PLGA was chemically treated in 0.1 N NaOH for various periods of time to transform traditionally nano-smooth PLGA into a nano-rough PLGA to improve osteoblast functions48 (Figure 4). It has also been shown that the promising nanosurface features present in the aforementioned metals and ceramics can be transferred to polymers to promote bone cell functions.49–51 In addition, Price et al.49 conducted a study of osteoblast behavior on polymer (PLGA) casts of consolidated carbon nanofiber-based materials, which have previously been shown to improve osteoblast functions compared to conventional carbon fiber materials.50 They demonstrated increased osteoblast adhesion on polymer casts

 2009 Jo h n Wiley & So n s, In c.

341

Advanced Review

(a)

www.wiley.com/wires/nanomed

(b)

FIGURE 4 | Scanning electron

micrographs of chemically treated PLGA surfaces. Representative scanning electron micrograph images of (a) chemically untreated (conventional) PLGA (feature dimensions 10,000–15,000 nm) and (b) chemically treated nano-structured PLGA (feature dimensions 50–100 nm). Scale bar = 100,000 nm. (Reprinted with permission from Ref 49. Copyright 2003 Elsevier).

of nanophase carbon fibers compared to polymer casts of conventional carbon fibers. Similarly, polymer casts of composites of polycarbonate urethane/CNTs also promoted osteoblast functions compared to casts of polycarbonate urethane/conventional carbon tubes.51 All this evidence once again signifies the important role of nanoscale surface structures for bone applications. Importantly, unlike other tissue engineering research areas, to create better polymeric orthopedic materials, the reasons why nanostructured polymers promote osteoblast functions are known. Specifically, like metals and ceramics, the mechanisms of increased adsorption of select proteins (such as fibronectin and vitronectin) important for osteoblast adhesion has been determined on nanophase polymers and can be correlated to the higher surface energy of such nanophase compared to conventional polymers.35,36 It is envisioned that one day nanotechnology will create a specific nanometer surface feature to generate surface energy values required to promote the adsorption of a select protein important for mediating the adhesion of a certain cell. This has already been partially observed for the use of nanostructured polymers for vascular, cartilage, and bladder applications where unique nanometer surface features are used for each type of tissue engineering application.38

Ceramic/Polymer Nanocomposites The previously reported promising results of nanophase ceramics and polymers for orthopedic applications independently led to the idea of combining the advantages of both types of materials to create even better bone scaffold materials.52–54 For example, scaffold materials (such as NHAP and other calcium phosphates) can facilitate greater scaffold strength and bone cell functions than conventional materials.55 However, by themselves HA and calcium phosphates have slow biodegradation rates. Therefore, to avoid this drawback biodegradable polymers can be incorporated into a composite to fabricate more favorable materials for orthopedic applications than the individual components.56 The more important 342

point is that the nanocomposites can mimic the constituents of natural bone to some extent better than individual components.57 Recent studies have suggested that better osteoconductivity can be achieved by synthetic composite materials that resemble the size and morphology of both the inorganic particles and organic phase of bone.58,59 Bone cell functions can be enhanced by interacting with nanophase ceramics and nanostructured polymers collectively compared to individually.60–64 For example, Jung et al. conducted research on PLA/calcium metaphosphate composites which showed that osteoblast functions on the composite were significantly enhanced compared to scaffolds made of PLA alone.65 The ceramic materials chosen for composite fabrication are often calcium phosphates (i.e., HA), metal oxides (i.e., titania, alumina), and glass ceramics (Bioglass). The choices for polymeric materials usually have been from the poly(α-hydroxy acid) family including PLA, PGA, and their copolymer PLGA; collagen has also often been chosen. For example, Du et al. developed nano-HA/collagen composite scaffolds which promoted the deposition of a new bone matrix at the interface of bone fragments and the composite. Furthermore, they also demonstrated that the porous nano-HA/collagen scaffold provided a microenvironment resembling that seen in vivo with osteoblasts within the composite eventually acquiring a three-dimensional polygonal shape.58 Nanocomposites of titania/PLGA were also prepared with different ultrasonic power to ensure dispersion of nanoceramic particles in the polymer; ceramic nanoparticle agglomeration in polymers is a common problem sometimes solved with sonication57,66 (Figure 5). The study demonstrated that osteoblast adhesion and long-term functions increased on nanocomposites prepared with greater sonication powers compared to that of scaffolds made of pure PLGA. Among those composites, the one that had the closest surface roughness values (as measured by atomic force microscopy) to natural bone showed the

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

(a)

(c)

(b)

(d)

FIGURE 5 | AFM images of nanophase titania and the PLGA mold of nanophase titania, conventional titania and the PLGA mold of conventional

titania. Images provided evidence of the successful transfer of the surface roughness of nanophase titania to PLGA molds of nanophase titania, and conventional titania to PLGA moulds of conventional titania. Root-mean-square values from AFM for nanophase titania at 5 rmmum × 5 µm and 25 µm × 25 µm scans were 29 and 22 nm, respectively. Root-mean-square values from AFM for the PLGA mold of nanophase titania at 5 µm × 5 µm and 25 µm × 25 µm scans were 35 and 27 nm, respectively. Root-mean-square values from AFM for conventional titania at 5 µm × 5 µm and 25 µm × 25 µm scans were 12 and 11 nm, respectively. Root-mean-square values from AFM for the PLGA mold of conventional titania at 5 µm × 5 µm and 25 µm × 25 µm scans were 13 and 12 nm, respectively. (a) Nanophase titania. (b) PLGA mold of nanophase titania. (c) Conventional titania. (d) PLGA mold of conventional titania. (Reprinted with permission from Ref 6. Copyright 2005 Institute of Physics Publishing (IOP)).

greatest osteoblast adhesion and subsequent calciumcontaining-mineral deposition.56 Specifically, up to 3 times more osteoblasts adhered to the nanophase titania/PLGA than the conventional titania/PLGA composites at the same weight ratio and porosity.66 Another advantage of nanocomposites compared to conventional materials is an improvement in mechanical properties. For instance, McManus et al. reported greater mechanical properties of polymer/ceramic composites with nanometer grain size compared to micron sized ones.67 Composites of PLA with 40 and 50 wt% nanophase (< 100 nm) alumina, titania, and HA showed significantly greater bending moduli (closer to bone) than that of composites with conventional coarser-grained ceramics. Specifically, the bending modulus of nanophase titania/PLA composites with a weight ratio 50/50 was 1960 ± 250 MPa, which was on the same order of magnitude of the trabecular bone.67 On the other hand, the bending modulus of plain PLA and conventional titania/PLA of the same weight ratio was only 60 ± 3 and 870 ± 30 MPa, respectively.67 Clearly, the Vo lu me 1, May /Ju n e 2009

enhanced mechanical and biocompatibility properties of nanostructured polymer/ceramic composites promise great improvements for orthopedic applications.

Magnetic Nanoparticles to Reverse Osteoporosis As a last example, in recent years nanotechnology has allowed for the possibility of fabricating nanophase magnetic particles with functions suitable for targeting and treating various bone diseases.68–71 In particular, iron oxide nanoparticles, which possess magnetic properties, are particularly promising for several orthopedic applications, such as72–76 (1) osteoblast cellular therapy including cell labeling and targeting and as a tool for cell-biology research to separate and purify osteoblast populations, (2) bone tissue repair, (3) bone drug delivery, (4) magnetic resonance imaging (MRI) of bone, (5) bone hyperthermia, (6) fighting bone cancer, (7) magnetofection. Materials with a high saturated magnetization (such as

 2009 Jo h n Wiley & So n s, In c.

343

Advanced Review

(a)

www.wiley.com/wires/nanomed

(b)

FIGURE 6 | Transmission electron microscope (TEM) images of (a) γ -Fe2 O3 and (b) Fe3 O4 magnetic nanoparticles. Scale bars = 20 nm.

transition metals (e.g., Fe, Co, Ni) or metal oxides (e.g., Fe3 O4 , γ -Fe2 O3 ) are usually considered for such magnetic nanoparticle applications. Although pure metals possess the highest magnetic saturation, they are highly toxic and extremely sensitive to oxidation,77 and therefore are not relevant for many biomedical applications inside the body. In contrast, iron oxides are less sensitive to oxidation and therefore can give stable magnetic responses. In fact, small iron oxide particles have been applied for in vitro diagnosis for around 50 years.78 Recent studies and investigations have showed that magnetite (Fe3 O4 ) and maghemite (γ -Fe2 O3 ) (Figure 6) are very promising candidates particularly for orthopedic applications owing to their biocompatibility and relative ease of functionalization.79–81 For practical purposes, these magnetic nanoparticles must not only combine the properties of high magnetic saturation and bio- and cytocompatibility, but they must also overcome a major difficulty when using nanoparticles: aggregation. Without any surface modification, magnetic iron oxide nanoparticles possess hydrophobic surfaces with large surface area-to-volume ratios.72 Because of the hydrophobic interactions between the particles, they agglomerate to form larger clusters, resulting in increased particle size. These agglomerations show strong dipole–dipole interactions and ferromagnetic behavior.82 Clusters will be further magnetized when in a magnetic field, causing stronger attractions between them83 and consequently creating increasing aggregation sizes. With a proper surface coating or functionalization with organic polymers or other inorganic metals (e.g., gold) or oxide surfaces (e.g., silica or alumina), these magnetic nanoparticles can be dispersed into suitable solvents.84 In all cases, magnetic nanoparticles in the size range of 10–100 nm are of interest because they exhibit superparamagnetic properties (meaning that they do not retain any magnetism after the removal of a magnetic field) and are small enough both to evade the reticuloendothelial clearance system of the body as well as to penetrate very small 344

capillaries/cells within the body tissues.72 Moreover, individual iron oxide nanoparticles can traverse the intricate porosity of bone to attach to specific regions of the diseased bone. Iron oxide magnetic nanoparticles can be prepared using several different methods based on two fundamental processes: size reduction and aqueous precipitation.77 Ball-mill grinding of micron-sized magnetite powders was one of the first methods developed by Papell in the late 1960s to form ∼10 nmdiameter magnetic particles.85 However, the requirement of 500–1000 h of grinding to form nanoparticles made this method inconvenient and it was largely replaced by aqueous precipitation methods. Other methods (such as gas-phase deposition and electronbeam lithography) are elaborate procedures unable to control the size of the magnetic particles.86–89 Aqueous precipitation methods for magnetite nanoparticle formation include oxidation of Fe2+ 90–92 and formation in water-in-oil microemulsions,93–97 vesicles,98 and apoferritin99–101 as well as in liposomes102,103 and in the presence of polymers.77,104–112 However, the most common synthetic route to fabricate iron oxides (e.g., Fe3 O4 and γ -Fe2 O3 ) is the coprecipitation of hydrated divalent and trivalent iron salts in the presence of a strong base.113–122 These methods are simple and more tractable and more efficient toward controlling magnetic particle size, composition, and sometimes even shape.72,123–125 For example, magnetite is usually prepared by adding a base (e.g., NaOH) to an aqueous mixture of Fe2+ and Fe3+ chloride at a 1 : 2 molar ratio. The precipitated magnetite is black in color. It should be noted that the reaction must be carried out under a nonoxidizing, oxygenfree environment. Otherwise, Fe3 O4 might be further oxidized to Fe(OH)3 . Studies showed that bubbling nitrogen through solution during magnetite synthesis can prevent further oxidation and also can reduce the particle size.125,126 The use of magnetic particles as drug delivery systems for nonorthopedic applications dates back to the 1970s142,127 . In the 1980s, many efforts were

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

made to develop delivery strategies using magnetic ¨ microcapsules and microspheres.128–131 Lubbe et al., were the first to use magnetic nanoparticles for animal model.132 in which small amounts of a ferrofluid were used as a vehicle to concentrate epirubicin locally in tumors. The studied later concluded that the ferrofluid did not cause any major laboratory abnormalities. Therefore, magnetic fluid may be a safe agent, which can be used in different ways for local forms of cancer treatment. The same group of scientists conducted a Phase I clinical trial in patients with advanced cancer and treated cancers using magnetic nanoparticles loaded with epirubicin.133 The results demonstrated that magnetic drug targeting with epirubicin was well tolerated and that the nanoparticles were successfully directed to the tumors in about one-half of the patients. More recently, using similar drug delivery processes, several groups have reported the successful cytotoxic delivery and tumor remission in animal models including swines,134,135 rabbits,136 and rats.137–139 Despite all these promising results in treating cancer research with magnetic nanoparticles, the use of iron oxide nanoparticles in orthopedic applications remains largely unexplored. Pareta et al. were one of the first to use magnetic nanoparticles for treating bone diseases (such as osteoporosis).140 Their general idea was to fabricate and modify magnetic nanoparticles with surfactants, drug coatings, and calcium phosphate coatings before their injection into a porous bone site. The coated magnetic nanoparticles would then be driven by an external magnetic field to specific areas of bone disease (such as osteoporotic bone). Eventually, after the magnetic field is turned off, the particles would attach to osteoporotic bone and immediately build bone mass simply by their presence at the site of weak bone. Specifically, in such studies, iron oxides nanoparticles (Fe3 O4 and γ -Fe2 O3 ) were prepared via wet chemistry methods under high pH as previously reported.141 All particles were magnetic, with sizes ranging from 10 to 20 nm in diameter. Bovine serum albumin (BSA) and citric acid (CA) were used as surfactants to prevent nanoparticle agglomeration. These nanoparticle solutions were later added into osteoblast cell culture media and incubated with the cells for various periods of time. The study showed for the first time significantly increased osteoblast density when cultured in the presence of BSA compared to controls (without any nanoparticles). Furthermore, the results also demonstrated greater osteoblast densities in the presence of maghemite nanoparticles than magnetite nanoparticles and controls after 5 and 8 days. While the mechanism of greater osteoblast growth Vo lu me 1, May /Ju n e 2009

in the presence of maghemite nanoparticles remains unclear, one proposed explanation was that after these time periods, the magnetite started to agglomerate in the cell culture media and hence decreased the effective surface area for cell interactions. Clearly, this study exhibited great potential in treating not only osteoporosis but also other local bone diseases (such as bone cancer) by tailoring such magnetic nanoparticles for that disease. Even though showing great potential in orthopedic applications, disease treatment with magnetic nanoparticles still has some limitations. The main limitation relates to the strength of the external magnetic field that must be applied to obtain the necessary gradient to control the nanoparticles in bone.142 A strong permanent NdFeB magnet can reach effective magnetic fields as deep as 15 cm in the body.143 However, depending on the geometry of the magnetic field, magnetic nanoparticles can be scattered around the desired site causing the delivery ineffective. Therefore, designing a proper magnetic field is extremely critical for magnetic nanoparticle targeting processes. Another limitation is that the superparamagnetic nanoparticles of small sizes possess much lower magnetic saturations compared to the bulk phase. Consequently, the drag force of blood flow and/or other physiological forces makes it more difficult to focus such nanoparticles to the bone target.77,142

FUTURE CHALLENGES Nanotechnology has developed to the extent that it is now possible to fabricate advanced materials with more favorable properties for orthopedic applications. However, it should be noted that studies on nanophase materials have only just begun; there are still many unanswered questions and unexplored frontiers which can greatly influence the role of nanostructured materials as improved bone implants.144,157 First, nanomaterial fabrication techniques can be improved with better resolution, accuracy, and expense; without inexpensive techniques that can precisely fabricate exact nanodimensions, industry may not participate in the use of nanomaterials in orthopedics. Second, more in vitro and in vivo investigations regarding biocompatibility and degradation behavior are needed. Lastly, more experiments need to be conducted to elucidate the potential toxicity of nanomaterials not only when in the body due to implantation but also when exposed to such particles during manufacturing. Regarding the use of magnetic nanoparticles for treating diseased bone, only a few studies even exist thus far. Toxicity is also a concern for these magnetic nanoparticles, which needs further investigation. It

 2009 Jo h n Wiley & So n s, In c.

345

Advanced Review

www.wiley.com/wires/nanomed

has been difficult to accurately report the toxicity of nanoparticles since toxicity depends on numerous factors including the dose, chemical composition, method of administration, size, biodegradability, solubility, pharmacokinetics, biodistribution, surface chemistry, shape, and structure, to name a few.142 It was suggested that the most important characteristics regarding cytotoxicity are size, shape, composition, and coating of nanoparticles.145 Among those, surface modification may be a key tool to minimize the toxicity of nanoparticles without imparting their function.146 Some studies have indicated that it is biologically unfavorable if nanoparticles are inhaled and absorbed via the lung (or swallowed and then absorbedacross

the gastrointestinal tract).147 Interestingly, another study demonstrated that at concentrations of 20–100 mg/ml, large magnetic particles showed higher cytotoxicity than smaller ones.148 Although there are still many challenges ahead, already some magnetic drug delivery systems (not for orthopedics) have been commercialized (such as MagNaGel,149,150 FluidMAG,151 and TargetMAG).151 In summary, with the knowledge gained so far, it appears that nanostructured metals, ceramics, polymers, and composites hold much promise for orthopedic applications: enough promise to warrant further exploration.

REFERENCES 1. Liu Q. Tissue engineering. In: Shi D, ed. Biomaterials and Tissue Engineering. Verlag Berlin Heidenberg: Springer; 2004, 219.

11. Webster TJ, Siegel RW, Bizios R. Enhanced functions of osteoblasts on nanophase ceramics. Biomaterials 2000, 21(17):1803–1810.

2. Smith R. Bone Health and Osteoporosis: A Report of the Surgeon General. Rockville, MD: US Department of Health and Human Services, Public Health Service, Office of the Surgeon General; 2004, 68–70.

12. Liu H, Webster TJ. Bioinspired nanocomposites for orthopedic applications. In: Webster TJ, ed. Nanotechnology for the Regeneration of Hard and Soft Tissues. Singapore: World Scientific; 2007, 3.

3. http://www.niams.nih.gov/Health Info/Bone/Osteo porosis/default.asp. Accessed July 2008. 4. Webster TJ. Nanophase ceramics as improved bone tissue engineering materials. Am Ceram Soc Bull 2003, 82(6):23–28. 5. Buser D, Nydegger T, Oxland T, Cochran DL, Schenk RK, et al.: Interface shear strength of titanium implants with a sanblasted and acid-etched surface: a biomechanical study in the maxilla of miniature pigs. J Biomed Mater Res 1999, 45:75. 6. Webster TJ. Nanophase ceramics: the future of orthopedic and dental implant material. In: Ying JY, ed. Nanostructured Materials. New York: Academy Press; 2001, 125–166. 7. Kaplan FS, Hayes WC, Keaveny TM, Boskey A, Einhorn TA. Biomaterials. In: Simon SP, ed. Orthopedic Basic Science. Columbus, OH: American Academy of Orthopedic Surgeons; 1994, 460–478. 8. Daniels AU, Chang MKO, Andriano KP. Mechanical properties of biodegradable polymers and composites proposed for internal fixation of bone. J Appl Biomater 1990, 1:57.

13. Wang M. Developing bioactive composite materials for tissue replacement. Biomaterials 2003, 24:2133–2151. 14. Small IA, Misiek D. A sixteen-year evaluation of the mandibular staple bone plate. J Oral Maxillofac Surg 1986, 44(60):60–66. 15. McKee GK. Total hip replacement—past, present and future. Biomaterials 1982, 3:130–135. 16. Steflik DE, Sisk AL, Parr GR, Gardner LK, Hanes PJ, Lake FT, Berkery DJ, Brewer P. Osteogenesis at the dental implant interface : High-voltage electron microscopic and conventional transmission electron microscopic observations. J Biomed Mater Res 1993, 27:791. 17. Black J. Biological Performance of Materials. New York: Marcel Dekker; 1981. 18. Sinibaldi K, Rosen H, Lin SK, De Angelis M. Tumors associated with metallic implants in animals. Clin Orthop Relat Res 1976, 118:257–266.

9. Terjesen T, Apalset K. The influence of different degrees of stiffness of fixation plates on experimental bone healing. J Orthop Res 1988, 6:292.

19. de Oliveira PT, Nanci A. Nanotexturing of titaniumbased surfaces upregulates expression of bone sialoprotein and osteopontin by cultured osteogenic cells. Biomaterials 2004, 25:403–413.

10. Webster TJ, Siegel RW, Bizios R. Osteoblast adhesion on nanophase ceramics. Biomaterials 1999, 20(13):1221–1227.

20. Webster TJ, Ejiofor JU. Increased osteoblast adhesion on nanophase metals Ti, Ti6Al4V, and CoCrMo. Biomaterials 2004, 25(19):4731–4739.

346

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

21. Yao C, Perla V, McKenzie JL, Slamovich EB, Webster TJ. Anodized Ti and Ti6AlV possessing nanometer surface features enhance osteoblast adhesion. J Biomed Nanotechnol 2005, 1(1):68–73.

35. Webster TJ, Schadler LS, Siegel RW, Bizios R. Mechanisms of enhanced osteoblast adhesion on nanophase alumina involve vitronectin. Tissue Eng 2001, 7:291–301.

22. Sirivisoot S, Yao C, Xiao X, Sheldon BW, Webster TJ. Greater osteoblast functions on multiwalled carbon nanotubes grown from anodized nanotubular titanium for orthopedic applications. Nanotechnology 2007, 18:365102.

36. Webster TJ, Siegel RW, Bizios R. Specific proteins mediate osteoblast adhesion on nanophase ceramics. J Biomed Mater Res 2000, 51:475.

23. Chung TW, Liu DZ, Wang SY, Wang SS. Enhancement of the growth of human endothelial cells by surface roughness at nanometer scale. Biomaterials 2003, 24:4655–4661. 24. Kunzler TP, Drobek T, Schuler M, Spencer ND. Systematic study of osteoblast and fibroblast response to roughness by means of surface morphology gradients. Biomaterials 2007, 28:2175–2182. 25. Zhao G, Raines AL, Wieland M, Schwartz Z, Boyan BD. Requirement for both micron- and submicron scale structure for synergistic responses of osteoblasts to substrate surface energy and topography. Biomaterials 2007, 28:2821–2829.

37. Agrawal CM, Ray RB. Biodegradable polymeric scaffolds for musculoskeletal tissue engineering. J Biomed Mater Res 2001, 55(2):141–150. 38. Balasundaram G, Webster TJ. An overview of nanopolymers for orthopedic applications. Macromol Biosci 2007, 7:635–642. 39. Rodrigues CVM, Serricella P, Linhares ABR, Guerdes RM, Borojevic R, et al.: Characterization of a bovine collagen-hydroxyapatite composite scaffold for bone tissue engineering. Biomaterials 2003, 24(27):4987–4997. 40. Dunn MG, Bellincampi LD, Tria AJ, Zawadsky JP. Preliminary development of a collagen-PLA composite for ACL reconstruction. J Appl Polym Sci 1997, 63(11):1423–1428.

26. Dyer MA, Ainslie KM, Pishko MV. Protein adhesion on siliconsupported hyperbranched poly(ethylene glycol) and poly(allylamine) thin films. Langmuir 2007, 23:7018–7023.

41. Liao SS, Cui FZ, Zhu Y. Osteoblast adherence and migration through three-dimensional porous mineralized collagen based composite: nHAC/PLA. J Bioact Compat Polym 2004, 19(2):117–130.

27. Khang D, Lu J, Yao C, Haberstroh KM, Webster TJ. The role of nanometer and sub-micron surface features on vascular and bone cell adhesion on titanium. Biomaterials 2008, 29:970.

42. Liu H, Webster TJ. Bioinspired nanocomposites for orthopedic applications. In: Webster TJ: ed. Nanotechnology for the Regeneration of Hard and Soft Tissues: Singapore: World Scientific; 2007, 24.

28. Hayashi H, Uchida A, Hamada H, Yoshikawa H, Shinto Y, et al.: Alumina ceramic prostheses for bone reactions to alumina oxide dental implants in man: a case report. Arch Orthop Trauma Surg 1992, 112:1–4.

43. Liu Q. Tissue engineering. In: Shi D, ed. Biomaterials and Tissue Engineering. Berlin/Heidenberg: Springer; 2004, 200.

29. Davies JE, Matsuda T. Extracellular matrix production by osteoblasts on bioactive substrate in vitro. Scanning Microsc 1988, 2(1445):52. 30. Sautier JM, Nefussi JR, Forest N. Ultrasonic study of bone formation on synthetic hydroxyapatite in osteoblasts cultures. Cell Mater 1991, 1:209–217. 31. Matsuda T, Davies JE. The in vitro response of osteoblasts to bioactive glass. Biomaterials 1987, 8:275–285. 32. LeGeros RZ: Calcium Phosphates in Oral Biology and Medicine. Basel: Karger; 1991. 33. Kokubo T, Ito S, Huang T, Sakka S, Kitsugi T, et al.: P-rich layer formed on high-strength bioactive glassceramic A-W. J Biomed Mater Res 1990, 24:331–343. 34. Nagano M, Nakamura T, Kokubo T, Tanahashi M, Ogawa M. Differences of bone bonding ability and degradation behavior in vivo between amorphous calcium phosphate and highly crystalline hydroxylapatite coating. Biomaterials 1996, 17:1771–1777.

Vo lu me 1, May /Ju n e 2009

44. Chen VJ, Ma PX. Nano-fibrous poly(L-lactic acid) scaffolds with interconnected spherical macropores. Biomaterials 2004, 25(11):2065–2073. 45. Karp JM, Shoichet MS, Davies JE. Bone formation on two-dimensional poly(DL-lactide-co-glycolide) (PLGA) films and three-dimensional PLGA tissue engineering scaffolds in vitro. J Biomed Mater Res A 2003, 64(2):388–396. 46. Liu X, Ma PX. Polymeric scaffolds for bone tissue engineering. Ann Biomed Eng 2004, 32(3):477–486. 47. Wei G, Ma PX. Structure and properties of nano-hydroxyapatite/polymer composite scaffolds for bone tissue engineering. Biomaterials 2004, 25(19):4749–4757. 48. Thapa A, Miller DC, Webster TJ. Nano-structured polymers enhance bladder smooth muscle cell function. Biomaterials 2003, 24:2915. 49. Price RL, Ellison K, Haberstroh KM, Webster TJ. Nanometer surface roughness increases select osteoblast adhesion on carbon nanofiber compacts. J Biomed Mater Res 2004, 70:129.

 2009 Jo h n Wiley & So n s, In c.

347

Advanced Review

www.wiley.com/wires/nanomed

50. Price RL, Gutwein LG, Kaledin L, Tepper F, Webster TJ. Osteoblast function on nanophase alumina materials: influence of chemistry, phase, and topography. J Biomed Mater Res 2003, 67:1284.

64. Ma PX, Zhang R, Xiao G, Franceschi R. Engineering new bone tissue in vitro on highly porous poly(ahydroxyl acids)/hydroxyapatite composite scaffolds. J Biomed Mater Res 2001, 54(2):284–293.

51. Webster TJ, Waid MC, McKenzie JL, Price RL, Ejiofor JU. Nano-biotechnology: carbon nanofibers as improved neural and orthopaedic implants. Nanotechnology 2004, 15:48.

65. Jung Y, Kim SS, Kim YH, Kim SH, Kim BS, et al.: A poly(lactic acid)/calcium metaphosphate composite for bone tissue engineering. Biomaterials 2005, 26:6314–6322.

52. Boccaccini AR, Maquet V. Bioresorbable and bioactive polymer/bioglass composites with tailored pore structure for tissue engineering applications. Compos Sci Technol 2003, 63:2417–2429.

66. Smith TA, Webster TJ. Increased osteoblast function on PLGA composites containing nanophase titania. J Biomed Mater Res 2005, 74A:677–686.

53. Hutmacher DW. Scaffolds in tissue engineering bone and cartilage. Biomaterials 2000, 21:2529–2543. 54. Thomson RC, Yaszemski MJ, Powers JM, Mikos AG. Hydroxyapatite fiber reinforced poly(αhydroxy ester) foams for bone regeneration. Biomaterials 1998, 19:1935–1943. 55. Kikuchi M, Itoh S, Ichinose S, Shinomiya K, Tanaka J. Self-organization mechanism in a bone-like hydroxyapaptite/collagen nanocomposite synthesized in vitro and its biological reaction in vivo. Biomaterials 2001, 22:1705–1711. 56. Christenson EM, Anseth KS, van den Beucken JJ, Chan CK, Ercan B, et al.: Nanobiomaterial applications in orthopedics. J Orthop Res 2007, 25(1):11–22. 57. Liu H, Slamovich EB, Webster TJ. Increased osteoblast functions on nanophase titania dispersed in poly-lactic-co-glycolic acid composites. Nanotechnology 2005, 16:S601–S608. 58. Du C, Cui FZ, Zhu XD, de Groot K. Threedimensional nano-HAP/collagen matrix loading with osteogenic cells in organ culture. J Biomed Mater Res 1999, 44:407–415. 59. Du C, Cui FZ, Feng QL, Zhu XD, de Groot K. Tissue response to nano-hydroxyapatite/collagen composite implants in marrow cavity. J Biomed Mater Res 1998, 42:540–548. 60. Boccaccini AR, Roether JA, Hench LL, Maquet V, Jerome R. A composite approach to tissue engineering. Ceram Eng Sci Proc 2002, 23(4):805–816. 61. Marra KG, Szem JW, Kumta PN, Di Milla PA, Weiss LE. In vitro analysis of biodegradable polymer blend/hydroxyapatite composites for bone tissue engineering. J Biomed Mater Res 1999, 47(3):324–335. 62. Kalita S, Finley J, Bose S, Hosick H, Bandyopadhyay A. Development of porous polymer-ceramic composites as bone grafts. Mater Res Soc Symp Proc 2002, 726:91–96. 63. Blaker JJ, Gough JE, Maquet V, Notingher I, Boccaccini AR. In vitro evaluation of novel bioactive composites based on bioglass—filled polylactide foams for bone tissue engineering scaffolds. J Biomed Mater Res A 2003, 67(4):1401–1411.

348

67. McManus AJ, Doremus RH, Siegel RW, Bizios R. Evaluation of cytocompatibility and bending modulus of nanoceramic/polymer composites. J Biomed Mater Res A 2005, 72(1):98–106. 68. Moghimi SM, Hunter ACH, Murray JC. Longcirculating and target-specific nanoparticles: theory to practice. Pharmacol Rev 2001, 53:283–318. 69. Curtis ASG, Wilkinson C. Nanotechniques and approaches in biotechnology. Trends Biotechnol 2001, 19:97–101. 70. Wilkinson JM. Nanotechnology applications in medicine. Med Device Technol 2003, 14(5):29–31. 71. Panyam J, Labhasetwar V. Biodegradable nanoparticles for drug and gene delivery to cells and tissue. Adv Drug Deliv Rev 2003, 55:329–347. 72. Gupta AK, Gupta M. Synthesis and surface engineering of iron oxide nanoparticles for biomedical applications. Biomaterials 2005, 26:3995–4021. 73. Arbab AS, Bashaw LA, Miller BR, Jordan EK, Lewis BK, et al.: Characterization of biophysical and metabolicproperties of cells labeled with superparamagnetic iron oxide nanoparticles and transfection agent for cellular MR imaging. Radiology 2003, 229(3):838–846. 74. Reimer P, Weissleder R. Development and experimental application of receptor-specific MR contrast media. Radiology 1996, 36:153–163. 75. Pankhurst QA, Connolly J, Jones SK, Dobson J. Applications of magnetic nanoparticles in biomedicine. J Phys D: Appl Phys 2003, 36:R167–R181. 76. Hafeli U, Schutt W, Teller J, Zborowski M, eds. Scientific and Clinical Applications of Magnetic Carriers. New York: Plenum Press; 1997. 77. Harris LA. Polymer Stabilized Magnetite Nanoparticles and Poly(propylene oxide) Modified Styrenedimethacrylate Networks. Doctor of Philosophy dissertation, Virginia Polytechnic Institute and State University, Virginia, 2002, 29. 78. Gilchrist RK, Medal R, Shorey WD, Hanselman RC, Parrot JC, et al.: Selective inductive heating of lymph nodes. Ann Surg 1957, 146:596–606.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

79. Schwertmann U, Cornell RM. Iron Oxides in the Laboratory: Preparation and Characterization. Weinheim: VCH; 1991.

95. O‘Connor CJ, Seip C, Sangregorio C, Carpenter E, Li S, Irvin G, John VT. Mol Cryst Liq Cryst 1999, 335:423–442.

80. Oliveira DM, Lacava ZGM, Lima ECD, Morais PC, Tedesco AC. Zinc phthalocyanine/magnetic fluid complex: a promising dual nanostructured system for cancer treatment. J Nanosci Nanotechnol 2006, 6:2432–2437.

96. Lopez-Perez JA, Lopez-Quintela MA, Mira J, Rivas J. Preparation of magnetic fluids with particles obtained in microemulsions. IEEE Trans Magn 1997, 33:4359.

¨ ¨ D, Kaufner L, Pison U, 81. Thunemann AF, Schutt ¨ Mohwald H. Maghemite nanoparticles protectively coated with poly(ethylene imine) and poly(ethylene oxide)-block-poly(glutamic acid). Langmuir 2006, 22:2351–2357. 82. Hamley IW. Nanotechnology with soft materials. Angew Chem Int Ed 2003, 42:1692–1712. 83. Tepper T, Ilievski F, Ross CA, Zaman TR, Ram RJ, et al.: Magneto-optical properties of iron oxide films. J Appl Phys 2003, 93(10):6948–6950. 84. Berry CC, Curtis ASG. Functionalisation of magnetic nanoparticles for applications in biomedicine. J Phys D: Appl Phys 2003, 36:R198–R206. 85. Papell SS, Faber OC. On the influence of nonuniform magnetic fields on ferro magnetic colloidal sols. NASA Technical Note, Vol. (NASA-TN-D-4676), 1968, 25. 86. Stolnik S, Illum L, Davis SS. Long circulating microparticulate drug carriers. Adv Drug Deliv Rev 1995, 16:195–214. 87. Kodas TT, Hampden-Smith M: Aerosol Processing of Materials. New York: Wiley-VCH; 1999. 88. Lee CS, Lee H, Westervelt RM. Microelectromagnets for the control of magnetic nanoparticles. Appl Phys Lett 2001, 79(20):3308–3310. 89. Rishton A, Lu Y, Altman RA, Marley AC, Bian Hahnes C, et al.: Magnetictunnel junctions fabricated at tenth-micron dimensions by electron beam lithography. Microelectron Eng 1997, 35:249. 90. Sidhu PS, Gilkes RJ, Posner AM. The synthesis and some properties of Co, Ni, Zn, Cu, Mn and Cd substituted magnetites. J Inorg Nucl Chem 1978, 40:429–437. 91. Strable E, Bulte JWM, Moskowitz B, Vivekanandan K, Allen M, et al.: Synthesis and characterization of soluble iron oxide-dendrimer composites. Chem Mater 2001, 13:2201. 92. Cornell RM, Schertmann U. Iron Oxides in the Laboratory; Preparation and Characterization.Weinheim: VCH; 1991. 93. Lee KM, Sorensen CM, Klabunde KJ, Hadjipanayis GC. Synthesis and characterization of stable colloidal Fe/sub 3/O/sub 4/ particles in water-in-oil microemulsions. IEEE Trans Magn 1992, 28:3180. 94. Dresco PA, Zaitsev VS, Gambino RJ, Chu B. Preparation and properties of magnetite and polymer magnetite nanoparticles. Langmuir 1999, 15:1945.

Vo lu me 1, May /Ju n e 2009

97. Santra S, Tapec R, Theodoropoulou N, Dobson J, Hebard R, Tan W. Synthesis and characterization of silica-coated iron oxide nanoparticles in microemulsion: The effect of nonionic surfactants. Langmuir 2001, 17:2900–2906. 98. Yaacob II, Nunes AC, Bose A. Magnetic nanoparticles produced in spontaneous cationic-anionic vesicles: room temperature synthesis and characterization. J Colloid Interface Sci 1995, 171:73. 99. Meldrum FC, Heywood BR, Mann S. Magnetoferritin: in vitro synthesis of a novel magnetic protein. Science 1992, 257:522. 100. Harrison PM, Arosio P. The ferritins: molecular properties, iron storage function and cellular regulation. Biochim Biophys Acta 1996, 1275:161. 101. Wong KKW, Douglas T, Gider S, Awschalom DD, Mann S. Biomimetic synthesis and characterization of magnetic proteins (magnetoferritin). Chem Mater 1998, 10:279. 102. Sangregorio C, Wiemann JK, O‘Connor CJ, Rosenzweig Z. A new method for the synthesis of magnetoliposomes. J Appl Phys 1999, 85:5699. 103. Bulte JWM, de Cuyper M, Despres D, Frank JA. Preparation, relaxometry, and bio kinetics of PEGylated magneto liposomes as MR contrast agent. J Magn Magn Mater 1999, 194:204–209. 104. Molday RS, MacKenzie D. Immuno specific ferro magnetic iron-dextron reagents for the labeling and magnetic separation of cells. J Immunol Methods 1982, 52:353–367. 105. Kim DK, Zhang Y, Kehr J, Klason T, Bjelke B, et al.. Characterization and MRI study of surfactant-coated superparamagnetic nanoparticles administered into the rat brain. J Magn Magn Mater 2001, 225:256. 106. Pardoe H, Chua-anusorn W, St. Pierre TG, Dobson J. Structural and magnetic properties of nanoscale iron oxide particles synthesized in the presence of dextran or polyvinyl alcohol. J Magn Magn Mater 2001, 225:41. 107. Lee J, Isobe T, Senna M. Preparation of ultrafine Fe3 O4 particles by precipitation in the presence of PVA at high pH. J Colloid Interface Sci 1996, 177:490. 108. Mendenhall GD, Geng Y, Hwang J. Optimization of long-term stability of magnetic fluids from magnetite and synthetic polyelectrolytes. J Colloid Interface Sci 1996, 184:519.

 2009 Jo h n Wiley & So n s, In c.

349

Advanced Review

www.wiley.com/wires/nanomed

109. Ding XB, Sun ZH, Wan GX, Jiang YY. Preparation of thermosensitive magnetic particles by dispersion polymerization. React Funct Polym 1998, 38:11. 110. Dumazet-Bonnamour I, Le Perchec P. Colloidal dispersion of magnetite nanoparticles via in situ preparation with sodium polyoxyalkylene di-phosphonates. Colloids Surf A 2000, 173:61. 111. Tamaura Y, Takahasi K, Kodera Y, Saito Y, Inada Y. Chemical modification of lipase with ferromagnetic modifier - A ferro magnetic - modified lipase. Biotechnol Lett 1986, 8:877–880. 112. Wormuth K. Superparamagnetic latex via inverse emulsion polymerization. J Colloid Interface Sci 2001, 241:366. 113. Blum E, Cebers A, Maiorov MM. Magnetic Fluids. Berlin: Walter de Gruyter; 1997. 114. Sato T, Iijima T, Sekin M, Inagaki N. Magnetic properties of ultrafine ferrite particles. J Magn Magn Mater 1987, 65:252–256. 115. Zins D, Cabuil V, Massart R. New aqueous magnetic fluids. J Mol Liq 1999, 83:217. 116. Goetze T, Gansau C, Buske N, Roeder M, Gornent P, Bahr M. Bio compatible magnetic core/shell nanoparticles. J Magn Magn Mater 2002, 252:399–402. 117. Elmore WC. Ferromagnetic colloid for studying magnetic structures. Phys Rev 1938, 54:309. 118. Bacri J, Perzynski R, Salin D, Cabuil V, Massart R. Ionic ferrofluids: a crossing of chemistry and physics. J Magn Magn Mater 1990, 85:27. 119. Massart R. Preparation of aqueous magnetic liquids in alkaline and acidic media. IEEE Trans Magn 1981, 17:1247. 120. Shen L, Laibinis P, Hatton TA. Bilayer surfactant stabilized magnetic fluids: synthesis and interactions at interfaces. Langmuir 1999, 15:447. 121. Gribanov NM, Bibik EE, Buzunov OV, Naumov VN. Physico-chemical regularities of obtaining highly dispersed magnetite by the method of chemical condensation. J Magn Magn Mater 1990, 85:7. 122. Wu KT, Kuo PC, Yao YD, Tsai EH. Magnetic and optical properties of Fe3 O4 nanoparticle ferrofluids prepared by coprecipitation techniques. IEEE Trans Magn 2001, 37:2651. 123. Gupta AK, Wells S. Surface modified superparamagnetic nanoparticles for drug delivery: preparation, characterization and cytotoxicity studies. IEEE Trans Nanobiosci 2004, 3(1):66–73. 124. Charles SW. Magneticfluids (ferrofluids). In: Dormann JL, Fiorani D, eds. Magnetic Properties of Fine Particles. North Holland: Elsevier; 1992, 267–374. 125. Gupta AK, Curtis ASG. Lactoferrin and ceruloplasmin derivatized superparamagneticiron oxide nanoparticles for targeting cell surface receptors. Biomaterials 2004, 25(15):3029–3040.

350

126. Kim DK, Zhang Y, Voit W, Rao KV, Muhammed M. Synthesis and characterization of surfactantcoated superparamagnetic monodispersed iron oxide nanoparticles. J Magn Magn Mater 2001, 225(1–2):30–36. 127. Zimmermann U, Pilwat G. Organ specific application of drugs by means of cellular capsule systems. Z Nat C Biosci 1976, 31:732–736. 128. Kato T, Nemoto R, Mori H, Abe R, Unno K, et al.: Magnetic microcapsules for targeted delivery of anticancer drugs. Appl Biochem Biotechnol 1984, 10:199. 129. Widder KJ, Morris RM, Poore G, Howard DP, Senyei AE. Tumor remission in yoshida sarcomabearing rats by selective targeting of magnetic albumin microspheres containing doxorubicin. Proc Natl Acad Sci USA 1981, 78:579. 130. Gupta PK, Hung CT. Effect of carrier dose on the multiple tissue disposition of doxorubicin hydrochloride administered via magnetic albumin microspheres in rats. J Pharmacol Sci 1989, 78:290. 131. Morimoto Y, Sughibayashi K, Okumura M, Kato Y. Biomedical applications of magnetic fluids. I. Magnetic guidance of ferro colloid entrapped albumin microspheres for site-specific drug delivery in vivo. J Pharmacobiodyn 1980, 3:264. ¨ 132. Lubbe AS, Bergemann C, Huhnt W, Fricke T, Riess H, et al.: Preclinical experiences with magnetic drug targeting: tolerance and efficacy. Cancer Res 1996, 56:4694. ¨ 133. Lubbe AS, Bergemann C, Riess H, Schriever F, Reichardt P, et al.: Clinical experiences with magnetic drug targeting: a phase I study with 4’-epidoxorubicin in 14 patients with advanced solid tumors. Cancer Res 1996, 56:4686. 134. Goodwin S, Peterson C, Hob C, Bittner C. Targeting and retention of magnetic targeted carriers (MTCs) enhancing intra-arterial chemotherapy. J Magn Magn Mater 1999, 194:132–139. 135. Goodwin SC, Bittner CA, Peterson CL, Wong G. Single-dose toxicity study of hepatic intra-arterial infusion of doxorubicin coupled to a novel magnetically targeted drug carrier. Toxicol Sci 2001, 60:177–183. 136. Alexiou C, Arnold W, Klein RJ, Parak FG, Hulin P, et al.: Locoregional cancer treatment with magnetic drug targeting. Cancer Res 2000, 60:6641–6648. 137. Lubbe AS, Bergemann C, Brock J, McClure DG. Physiological aspects in magnetic drug-targeting. J Magn Magn Mater 1999, 194:149–155. 138. Pulfer SK, Gallo JM. Enhanced brain tumor selectivity of cationic magnetic polysaccharide microspheres. J Drug Target 1999, 6:215–228. 139. Pulfer SK, Ciccotto SL, Gallo JM. Distribution of small magnetic particles in brain tumor-bearing rats. J Neurooncol 1999, 41:99–105.

 2009 Jo h n Wiley & So n s, In c.

Vo lu me 1, May /Ju n e 2009

WIREs Nanomedicine and Nanobiotechnology

Nanotechnology for bone materials

140. Pareta R, Taylor E, Webster TJ. Increased osteoblast density in the presence of novel calcium phosphate coated magnetic nanoparticles. Nanotechnology 2008, 19:7. 141. Kang YS, Risbud S, Rabolt JF, Stroeve P. Synthesis and characterization of nanometer-size Fe3 O4 and γ -Fe2 O3 particles. Chem Mater 1996, 8:2209–2211. ´ 142. Arruebo M, Fernandez-Pacheco R, Ibarra MR, Santamar´ıa J. Magnetic nanoparticles for drug delivery. Nano Today 2007, 2(3):22–32. ¨ B, Hofmann H, Hofmann M, 143. Neuberger T, Schopf von Rechenberg B. Superparamagnetic nanoparticles for biomedical applications: possibilities and limitations of a new drug delivery system. J Magn Magn Mater 2005, 293:483. 144. Liu H, Webster TJ. Bioinspired nanocomposites for orthopedic applications. In: Webster TJ, ed. Nanotechnology for the Regeneration of Hard and Soft Tissues: World Scientific; 2007, 43. 145. Macaroff PP, Simioni AR, Lacava ZGM, Lima ECD, Morais PC, et al.: Studies of cell toxicity and binding of magnetic nanoparticles with blood stream macromolecules. J Appl Phys 2006, 99:08S102. 146. Park SI, Lim JH, Kim JH, Yun HI, Kim CO. Toxicity estimation of magnetic fluids in a biological test. J Magn Magn Mater 2006, 304:e406. 147. Duncan R, Izzo L. Dendrimer biocompatibility and toxicity. Adv Drug Deliv Rev 2005, 57:2215.

148. Yin H, Too HP, Chow GM. The effects of particle size and surface coating on the cytotoxicity of nickel ferrite. Biomaterials 2005, 26:5818. 149. Sunderland CJ, Steiert M, Talmadge JE, Derfus AM, Barry SE. Targeted nanoparticles for detecting and treating cancer. Drug Dev Res 2006, 67:70. 150. www.alnis.com (accessed, July 2008). 151. www.chemicell.com (accessed, July 2008). 152. Liu H, Webster TJ. Bio inspired nano composites for orthopedic applications. In: Webster TJ, ed. Nanotechnology for the Regeneration of Hard and Soft Tissues. World Scientific, Singapore; 2007, 4. 153. Liu Q. Tissue engineering. In: Shi D, ed. Biomaterials and Tissue Engineering. Berlin Heidenberg: Springer Verlag; 2004, 208. 154. Liu Q. Tissue engineering. In: Shi D, ed. Biomaterials and Tissue Engineering. Berlin Heidenberg: Springer Verlag; 2004, 208. 155. Ma PX. Scaffolds for tissue fabrication. Materials Today 2004, 7(5):30–40. 156. Taddei P, Monti P, Simoni R. Vibrational and thermal study on the in vitro and in vivo degradation of a poly (lactic acid) - based bioabsorbable periodontal membrane. Journal of Materials Science : Materials in Medicine 2002, 13(5):469–475. 157. Balasundaram G. Nanomaterials for better orthopedics. In: Webster TJ, ed. Nanotechnology for the Regeneration of Hard and Soft Tissues. World Scientific, Singapore, 2007, 72.

FURTHER READING Palin E, Liu H, Webster TJ. Mimicking the nanofeatures of bone increases bone-forming cell adhesion and proliferation. Nanotechnology 2005, 16:1828–1835.

Vo lu me 1, May /Ju n e 2009

 2009 Jo h n Wiley & So n s, In c.

351