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Associate editor: M. Endoh

Current state, achievements, and future prospects of polymeric micelles as

nanocarriers for drug and gene delivery

Nobuhiro Nishiyama a, c, Kazunori Kataoka a,b,c,⁎

a Center for Disease Biology and Integrative Medicine, Graduate School of Medicine, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-0033, Japan b Department of Materials Engineering, Graduate School of Engineering, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-8656, Japan

c Center for NanoBio Integration, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-8656, Japan

Abstract

Polymeric micelles, self-assemblies of block copolymers, are promising nanocarrier systems for drug and gene delivery. Until now, several

micellar formulations of antitumor drugs have been intensively studied in preclinical and clinical trials, and their utility has been demonstrated.

Even compared with long-circulating liposomes, polymeric micelles might have several advantages, such as controlled drug release, tissue-

penetrating ability and reduced toxicity such as hand–foot syndrome and hypersensitivity reaction. Importantly, critical features of the polymeric

micelles as drug carriers, including particle size, stability, and loading capacity and release kinetics of drugs, can be modulated by the structures

and physicochemical properties of the constituent block copolymers. Also, nano-engineering of block copolymers might allow the preparation of

polymeric micelles with integrated smart functions, such as specific-tissue targetability, as well as chemical or physical stimuli-sensitivity. Thus,

polymeric micelles are nanotechnology-based carrier systems that might exert the activity of potent bioactive compounds in a site-directed

manner, ensuring their effectiveness and safety in the clinical use.

© 2006 Elsevier Inc. All rights reserved.

Keywords: Nanotechnology; Polymeric micelles; Cancer targeted therapy; Drug delivery; Gene and siRNA delivery; Photodynamic therapy

Abbreviations: AMD, age-related macular degeneration; ASGP, asialoglycoprotein; AUC, area under the curve; bFGF, basic fibroblast growth factor; c.a.c., critical

association concentration; CaP, calcium phosphate; CDDP, cisplatin [cis-dichlorodiammineplatinum (II)]; CNV, choroidal neovascularization; CP4, quadruplicated

cationic peptide; DACHPt, 1,2-diaminocyclohexane platinum (II); Dox, doxorubicin; Doxil, liposomal formulation of doxorubicin; DP, dendritic porphyrin; DPc,

dendritic phthalocyanine; EC, endothelial cells; ECM, extracellular matrix; EPR effect, enhanced permeability and retention effect; FBP, folate-binding proteins;

IC50, 50% growth inhibitory concentration; NCC, National Cancer Center; NK105, micellar formulation of paclitaxel; NK911, micellar formulation of doxorubicin;

oxaliplatin, oxalate 1,2-diaminocyclohexane platinum(II); NLS, nuclear localization signal; PCI, photochemical internalization; pDNA, plasmid DNA; PDT,

photodynamic therapy; PEG, poly(ethylene glycol); PEG-b-P(Asp), PEG-block -poly(aspartic acid); PEG-b-PDLLA, PEG-block -poly(D,L-lactide); PEG-b-P(Glu),

PEG-block -(glutamic acid); PEG-b-PLL, PEG-block -poly(L-lysine); PEG-liposome, PEG-modified liposome; PEI, polyethylenimine; PHPMA, N -(2-hydroxypropyl)

methacrylamide copolymer; PIC micelle, polyion complex micelle; PMPA, poly[(3-morpholinopropyl) aspartamide]; PS, photosensitizer; PTX, paclitaxel; RES,

reticuloendothelial system; siRNA, small interfering RNA; TEM, transmission electron microscopy; VEGF, vascular endothelial growth factor; Visudyne, liposomal

formulation of verteporfin; VPF, vascular permeability factor; VVO, vesicular vacuolar organelle.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 631

2. Biological significance of polymeric micelles. . . . . . . . . . . . . . . . . . . . . 632

2.1. Biodistribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 632

2.2. Accumulation in solid tumors . . . . . . . . . . . . . . . . . . . . . . . . . 632

3. Polymeric micelles for cancer chemotherapy . . . . . . . . . . . . . . . . . . . . . 633

Pharmacology & Therapeutics 112 (2006) 630–648

www.elsevier.com/locate/pharmthera

⁎ Corresponding author. Department of Materials Engineering, Graduate School of Engineering, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-

8656, Japan.

E-mail address: [email protected] (K. Kataoka).

0163-7258/$ - see front matter © 2006 Elsevier Inc. All rights reserved.doi:10.1016/j.pharmthera.2006.05.006

mailto:[email protected]

http://dx.doi.org/10.1016/j.pharmthera.2006.05.006

http://dx.doi.org/10.1016/j.pharmthera.2006.05.006

mailto:[email protected]



3.1. Polymeric micelles for delivery of hydrophobic drugs . . . . . . . . . . . . 633

3.2. Comparison between polymeric micelles and PEG-modified liposomes . . . 634

3.3. Polymer-metal complex micelles for delivery of platinous drugs . . . . . . . 635

3.4. Smart polymeric micelles for site-specific drug delivery . . . . . . . . . . . 637

3.5. Effects of polymeric nanocarriers on intracellular drug action . . . . . . . . 638

4. Dendritic photosensitizer-loaded polymeric micelles for photodynamic therapy . . . 639

5. Nanodevices for gene therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6415.1. Polyion complex (PIC) micelles for plasmid DNA delivery . . . . . . . . . 641

5.2. PIC micelles for small interfering RNA delivery . . . . . . . . . . . . . . . 642

5.3. Novel gene carriers enveloped in dendritic photosensitizer for

l ight-induced gene transfer . . . . . . . . . . . . . . . . . . . . . . . . . . 644

6. Future prospects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 645

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 645

1. Introduction

Recently, medical applications of nanotechnology haveattracted growing interest. Until now, a large number of new

nanotechnology-based concepts for therapeutic and diagnostic

medicines have emerged, and their feasibility has been

demonstrated (Ferrari, 2005). In 2005, the National Cancer

Institute (NCI) started the Cancer Nanotechnology Plan

(CNPlan) as a 10-year project to develop nanotechnologies,

which radically change the method of treatment, diagnosis and

prevention of cancers. In particular, considerable attention is

being focused on the nanotechnology-based drug delivery

because its strategic rationale has been demonstrated. During

the past decade, polymeric drug carriers including polymer –

drug conjugates and polymeric micelles have proven to be

useful in drug delivery, and several formulations have beenstudied in clinical trials (Duncan, 2003). Especially, polymeric

micelles are currently recognized as one of the most promising

modalities of drug carriers (Allen et al., 1999; Kataoka et al.,

1993, 2001; Lavasanifar et al., 2002; Torchilin, 2002;

Nishiyama et al., 2005b; Nishiyama & Kataoka, 2006).

The formation and characteristic properties of polymeric

micelles are illustrated in Fig. 1. It is well known that block

copolymers with amphiphilic character spontaneously assembleinto polymeric micelles with a diameter of several tens of

nanometers in aqueous media. Polymeric micelles have a

unique core–shell structure, in which an inner core serving as a

nanocontainer of hydrophobic drugs is surrounded by an outer

shell of hydrophilic polymers, such as poly(ethylene glycol)

(PEG), and have demonstrated longevity in the bloodstream and

effective tumor accumulation after their systemic administration

(Kwon et al., 1994; Yokoyama et al., 1999; Nishiyama et al.,

2003d; Bae et al., 2005b). Also, polymeric micelles have

several advantages, such as a simple preparation, efficient drug

loading without chemical modification of the parent drug, and

controlled drug release (Kataoka et al., 2001; Lavasanifar et al.,

2002; Nishiyama et al., 2005b; Nishiyama & Kataoka, 2006).Besides hydrophobic interaction, electrostatic interaction

between charged block copolymers and oppositely charged

macromolecules has allowed the formation of core–shell

nanoparticles, which are termed “ polyion complex (PIC)

micelles” (Harada & Kataoka, 1995, 1999; Kataoka et al.,

Fig. 1. Polymeric micelles as intelligent nanocarriers for drug and gene delivery.

631 N. Nishiyama, K. Kataoka / Pharmacology & Therapeutics 112 (20 06) 630 – 648



2001). This system is potentially useful for the delivery of genes

and small interfering RNA (siRNA) (Kataoka et al., 1996;

Katayose & Kataoka, 1997). It is known that the lack of

appropriate carriers for genes and siRNA remains a serious

problem for clinical application (Verma & Somia, 1997; Pack et

al., 2005). Also, recent advances in synthetic polymer chemistry

and biotechnology have allowed the development of polymericmicelles with integrated smart functions, such as environmen-

tally sensitivity and specific tissue targetability ( Nishiyama et

al., 2005b; Nishiyama & Kataoka, 2006). Thus, there is a strong

impetus for the development of polymeric micelle nanocarriers

to achieve successful drug and gene delivery.

This paper reviews recent progress in research on polymeric

micelles as nanocarriers for drug delivery, summarizing

accomplishments mainly of our group. Focus is placed on the

biological significance and future prospects of the polymeric

micelle-based nanocarriers.

2. Biological significance of polymeric micelles

2.1. Biodistribution

A major objective of using polymeric micelles as a drug

vehicle is to modulate drug disposition in the body directed

toward better therapeutic efficacy. For successful drug

targeting, the achievement of a prolonged blood circulation

of polymeric nanocarriers might be of primary importance,

because polymeric carriers are delivered to the target tissue

through the bloodstream, and the extravasation process is

generally considered to be slow and in a passive manner.

However, there are several obstacles to the long circulation of

polymeric carriers, which include glomerular excretion by thekidney and recognition by the reticuloendothelial system

(RES) located in the liver, spleen and lung (Kataoka, 1996).

The glomerular excretion can be avoided by using polymeric

carriers with a larger size than its threshold value (42–50 kDa

for water-soluble polymers; Seymour et al., 1987). On the

other hand, RES recognition may be avoidable by designing

polymeric carriers to have a size smaller than 200 nm as well

as an excellent biocompatibility (Stolnik et al., 1995;

Mosqueria et al., 2001). It is known that non-biocompatible

nanoparticles are recognized by the RES via the complement

activation, followed by elimination from the circulation;

however, the surface modification of nanoparticles withhydrophilic and biocompatible polymers, such as PEG, can

impair or even avoid RES recognition (Stolnik et al., 1995;

Mosqueria et al., 2001). A highly flexible and hydrated PEG

chain attached to the nanoparticle surface is assumed to have

an effective protein-resistant property due to its steric

repulsion effect (Jeon et al., 1991). Therefore, it is likely

that polymeric micelles, nanoscale colloidal carriers covered

with a high density of PEG shells, might circumvent the

aforementioned obstacles, thus showing a stealth property

during blood circulation.

Polymeric micelles are characterized by a critical association

concentration (c.a.c.), which defines a threshold concentration

for assembly. Polymeric micelles may not necessarily dissociate

immediately after extreme dilution following intravenous

injection into the body because they have a remarkably low

c.a.c. (10–6 – 10–7 M), which is 1000-folds lower than that of

surfactant micelles (La et al., 1996; Yamamoto et al., 2002), and

their dissociation is kinetically slow. This property allows the

micelles to circulate in the bloodstream until accumulation at

target tissues. The typical biodistribution of polymeric micellescan be exemplified by the results of polymeric micelles

composed of PEG-block -poly(D,L-lactide) (PEG-b-PDLLA)

copolymers labeled with 125I (Yamamoto et al., 2001). The

PEG-b-PDLLA micelles showed a remarkably prolonged blood

circulation (t 1/2∼18 hr) after intravenous administration, and

maintained 25% of the injected dose in the circulation at 24-hr

post-injection. The distribution volume in the central compart-

ment (V 0) and plasma-to-blood ratio of the micelles were

calculated to be nearly equivalent to the blood volume and the

plasma space value (= 1–hematocrit), respectively, suggesting

that polymeric micelles might distribute only to the blood

compartment and hardly interact with blood cells immediatelyafter their administration. Such a stable circulation of polymeric

micelles was also confirmed by a gel chromatography assay

(Yamamoto et al., 2001). Regarding tissue distribution, the

PEG-b-PDLLA micelles showed tissue-to-blood concentration

ratios ( K b) comparable to vascular space volume (∼0.2) in

normal organs (lung, kidney, liver and spleen) until 24 hr after

injection, and thereafter minimal increases in the liver and

spleen corresponding to an extracellular space volume (∼0.3)

(Yamamoto et al., 2001). Such low K b values of the polymeric

micelle in the liver and spleen were comparable to those

obtained from long-circulating liposomes (Allen & Hansen,

1991; Woodle et al., 1992). Thus, polymeric micelle avoided the

RES recognition as well as the entrapment by hepatic sinusoidalcapillaries characterized by large interendothelial junctions

(∼100 nm) and the absence of the basement membranes,

despite the relatively small size (∼30 nm) of the micelles.

Furthermore, it was revealed that the constituent block

copolymers might be finally excreted into the urine due to

their molecular weight being lower than the threshold of

glomerular filtration, suggesting the safety of polymeric

micelles with a low risk of chronic accumulation in the body.

2.2. Accumulation in solid tumors

It has been demonstrated that long-circulating polymericcarriers can preferentially and effectively accumulate in solid

tumors. This phenomenon is explained by the microvascular

hyperpermeability to circulating macromolecules and their

impaired lymphatic drainage in solid tumors, and is termed

the “Enhanced Permeability and Retention (EPR) effect ”

(Matsumura & Maeda, 1986; Maeda, 2001). Such tumor

vascular hyperpermeability has been suggested to be due to

overexpression of vascular permeability factor (VPF)/vascular

endothelial growth factor (VEGF) (Dvorak et al., 1995), as well

as secretion of other factors, such as the basic fibroblast growth

factor (bFGF) (Jain, 2001), bradykinin, nitric oxide and

peroxynitrate in tumor tissues (Maeda, 2001). Particularly,

VPF/VEGF, a tumor-secreted protein, may play active roles in

632 N. Nishiyama, K. Kataoka / Pharmacology & Therapeutics 112 (2006) 630 – 648



the angiogenesis process including vascular endothelial cell

(EC) division, selective degradation of vascular basement

membrane and surrounding extracellular matrix (ECM), and

EC migration as well as increased microvascular permeability

(Dvorak et al., 1995). Indeed, decreased vascular permeability

in colon carcinomas was observed when treated with the anti-

VEGF antibody (Yuan et al., 1996). Concerning the transvas-cular transport pathways of macromolecules, the involvement

of interendothelial junctions, transendothelial channels, fenes-

trations, and vesicular vacuolar organelles (VVO) have been

suggested by the morphological studies (Kohn et al., 1992; Neal

& Michel, 1995; Roberts & Palade, 1997; Hobbs et al., 1998 ),

but they remain controversial. Transmission electron micros-

copy (TEM) observations revealed that long-circulating lipo-

somes extravasate through either interendothelial or

transendothelial open junctions (Hobbs et al., 1998). Dvorak

et al. suggested that VVO, which are grape-like clusters of

vesicles and vacuoles with a size of 60–80 nm overlying the

entire thickness of EC cytoplasm from the luminal to theabluminal plasma membranes are probably responsible for the

transendothelial transport of the macromolecules (Kohn et al.,

1992; Dvorak et al., 1995).

To date, there is firm evidence that polymeric carriers

accumulate in various types of tumors, which is most probably

due to the above-mentioned EPR effect. The EPR effect appears

to be a phenomenon universally observed in malignant tumors.

Jain et al. reported the vascular cut-off sizes (the upper limit of

the size of long-circulating liposomes or latex beads which can

extravasate) ranging between 380 and 780 nm for 1 human and

5 murine tumors including mammary and colorectal carcino-

mas, hepatoma, glioma, and sarcoma (Hobbs et al., 1998; Jain,

2001). Hence, the vascular pore cut-off sizes of tumors areunlikely to be a significant obstacle to transvascular transport of

polymeric carriers with a relatively small size (i.e., less than

100 nm). Indeed, it was demonstrated that polymeric micelle

nanocarriers show an enhanced accumulation in solid tumors

(Kwon et al., 1994; Yokoyama et al., 1999; Nishiyama et al.,

2003d; Hamaguchi et al., 2005). The EPR effect is a strategic

basis for designing polymeric carriers for successful tumor-

targeted therapy.

3. Polymeric micelles for cancer chemotherapy

3.1. Polymeric micelles for delivery of hydrophobic drugs

Amphiphilic block copolymers spontaneously form core–

shell type polymeric micelles in aqueous media. Polymeric

micelles have a solid-like inner core, which serves as a potent

nanocontainer of hydrophobic compounds. The chemical

structures and properties of the micellar core-forming blocks

significantly affect drug loading efficiency and drug release

rate.

PEG-block -poly(aspartic acid) [PEG-b-P(Asp)] copolymers

chemically conjugated with doxorubicin (Dox) spontaneously

form polymeric micelles with a diameter of 15–60 nm. This

polymeric micelle can efficiently entrap free Dox in the inner

core, and the optimized formulation called NK911 is now being

studied in a phase II clinical trial at the National Cancer Center

(NCC) Hospital in Japan (Yokoyama et al., 1989, 1990a, 1990b,

1991, 1993; Kwon et al., 1994; Yokoyama et al., 1994, 1999;

Matsumura et al., 2004). In this formulation, Dox chemically

conjugated to the polymer side chain is pharmacologically

inactive, but contributes to the stable physical entrapment of

free Dox into the micellar core through π–π interaction of theanthracycline structure in Dox between the conjugated and

unconjugated ones, also allowing its sustained release from the

micellar core (Yokoyama et al., 1993, 1994). Thus, compati-

bility between the core-forming blocks and the drugs to be

loaded may be one of the most important factors increasing the

drug loading capacity as well as controlling the drug release

rate. Indeed, it was reported that very hydrophobic compounds

such as amphotericin B and the spicamycin derivative having a

long-chain fatty acid (KRN5500) are successfully incorporated

into polymeric micelles from PEG-b-P(Asp) derivatives having

fatty acid esters in the side chain (Yokoyama et al., 1998;

Lavasanifar et al., 2000). In addition to such structure matching between the inner core-forming blocks and drugs, the drug

loading amount should also be taken into consideration for

controlling the drug release kinetics. It was reported that a high

drug loading of the polymeric micelles resulted in the crystalline

formation of the loaded drugs inside the micellar core, leading

to a slower drug release rate (from 14 hr to 5 days) (Gref et al.,

1994; Jeong et al., 1998).

Due to the characteristic structure of the polymeric micelles,

of which the inner core is segregated by dense PEG palisade, the

properties or amount of the loaded drugs hardly affect their

biodistribution after systemic administration. Recently, pacli-

taxel (PTX) was incorporated into the polymeric micelles from

the PEG-b-P(Asp) modified with 4-phenyl-1-butanolate, andthis formulation was termed NK105 (Hamaguchi et al., 2005).

Figs. 2A and B show the plasma and tumor concentration of

PTX, respectively, after intravenous injection of free PTX or

NK105. In these figures, NK105 achieved approximately 90-

and 25-folds higher plasma and tumor areas under the curve

(AUC), respectively, than free PTX. Consequently, NK105

showed a remarkably enhanced antitumor activity against a

human colorectal cancer HT-29 cell xenograft in mice compared

to free PTX, while significantly restraining neurotoxicity, a

dose-limiting factor of PTX. The formulation of NK105 is

currently being studied in a phase I clinical trial at the NCC in

Japan.Recently, NK911, a micellar formulation of Dox was applied

to prevent the neointimal formation after balloon injury in the

rat carotid artery (Uwatoku et al., 2003). The balloon injury

induced a marked and sustained increase in the vascular

permeability, thereby allowing the effective accumulation of

NK911 in the balloon-injured artery. As a result, 3 intravenous

injections of NK911 significantly inhibited the neointimal

formation at 4 weeks after the balloon injury in not only single

but also double injury models (Fig. 3). In contrast, the treatment

with free Dox had no significant effects on the inhibition of the

neointimal formation. It was also found that the NK911

treatment induced neither the expression of several cytokines

nor systemic side effects. These results might offer novel




applications of polymeric nanocarriers for the prevention of

restenosis after balloon angioplasty and coronary stenting.

Thus, the enhanced accumulation of polymeric nanocarriers

might occur not only in solid tumors but also in other diseased

sites. Similarly, it was reported that polymeric micelles

accumulated in the area of the experimental myocardial

infarction (Lukyanov et al., 2004). More recently, we reportedthat polymeric micelles accumulated in the choroidal neovas-

cularization (CNV) sites in rat eyes (Ideta et al., 2005), offering

a new nanotechnology-based treatment of ophthalmic neovas-

cular diseases (see Section 4).

3.2. Comparison between polymeric

micelles and PEG-modified liposomes

PEG-modified liposomes (PEG-liposomes), which are

called “long-circulating liposomes” or “stealth liposomes”,

have been widely used as drug carriers for systemic injection.

In general, PEG-liposomes show a longer blood circulation period (t 1/2> 48 hr) than that of polymeric micelles (t 1/2< 24 hr).

Also, some liposomal formulations, such as Doxil® (Alza Co.)

and Visudyne® (Novartis Co.), have already been approved for

clinical use. Nevertheless, polymeric micelles have recently

attracted much attention due to their prominent properties over

those of the PEG-liposomes.

It has been suggested that the treatment with Doxil

sometimes induces the side effects of the hand–foot syndrome

as well as infusion-related reactions; therefore, the patients

need to be pretreated with anti-histamine or anti-inflammatory

agents before the administration of Doxil (Uziely et al., 1995;

Muggia et al., 1996; Stewart et al., 1998; Gordon et al., 2001).

Such a toxicity of Doxil appears to be a general problem in theclinical use of PEG-liposome formulations. The micellar

Fig. 3. Inhibitory effects of NK911 or Dox (3 times injection) on the neointimal formation in the rat carotid artery at 4 weeks after the balloon injury (single-injurymodel). The tissue sections were stained with H&E staining (Copyright 2003 American Heart Association).

Fig. 2. PTX concentration in the plasma (A) and tumor (B) after intravenous

injectionof NK105 orfreePTX to Colon 26-bearing mice (PTX dose: 100mg/kg)

(Hamaguchi et al., 2005).




formulations of Dox (NK911) and PTX (NK105) are currently

being studied in phase II and phase I clin ical trials,

respectively. Although the number of the patients treated

with these micellar drugs is much lower than that of the

patients treated with Doxil, such side effects encountered with

the use of Doxil have not been observed in the clinical trials of

NK911 and NK105 (Matsumura et al., 2004). It is assumedthat polymeric micelles composed of synthetic polymers might

not cause unfavorable biological responses inducible by

biomolecules such as lipids.

In regard to the therapeutic efficacy, Doxil is clinically

effective against ovarian cancer and breast cancer, both of

which feature a high density of tumor microvessels (Muggia,

2001). However, Doxil is not effective against stomach cancer

and pancreatic cancer. Probably, PEG-liposomes with a size of

ca. 150 nm might show limited tissue penetration in solid

tumors. It was reported that PEG-liposomes were mainly

localized outside tumor vessels, almost all around the vessel

wall, even 2 days after the systemic injection (Unezaki et al.,1996). In contrast, polymeric micelles having a smaller particle

size than that of PEG-liposomes are expected to show a higher

tumor-infiltrating ability. Indeed, it was demonstrated that

polymeric micelles with a size of 65 nm could access the inside

of 200-μm multicellular tumor spheroids (Bae et al., 2005a),

whereas the PEG-liposomes were localized on the periphery of

the spheroids (Tsukioka et al., 2002) (see Section 3.4). In

addition to such a limited tissue penetration, PEG-liposomes

have difficulties in incorporating hydrophobic drugs, which

easily form aggregates. It is assumed that the incorporation of a

large amount of such hydrophobic compounds might impair the

integrity of the liposomal membranes composed of lipid

bilayers. Also, PEG-liposomes have been suggested to be toostable to release loaded Dox, thus leading to a significant

reduction in the antitumor activity. In contrast, polymeric

micelles might allow the incorporation of various drugs

including very hydrophobic drugs, a high drug loading capacity

as well as the drug release in a sustained or site-specific manner.

These critical parameters can be controlled by optimization of

the chemical structure of the constituent block copolymers,

which is a strong motivation to develop polymeric micelle

nanocarriers.

3.3. Polymer-metal complex

micelles for delivery of platinous drugs

Cisplatin [cis-dichlorodiammineplatinum (II)] (CDDP) (Fig.

4A) is a metal complex antitumor agent widely used for the

treatment of many malignancies (Rosenberg, 1978). However,

its clinical use is limited due to toxic side effects such as acute

nephrotoxicity and chronic neurotoxicity. Also, it is known that

CDDP exhibits a very short circulation period after its systemic

injection while a significant amount undergoes glomerular

excretion (Siddik et al., 1987). Despite the optimization of the

dose regimen, the development of tumor-targetable formula-

tions of CDDP is demanded to improve the therapeutic efficacy

as well as the quality of life of the patients (Gianasi et al., 1999;

Newman et al., 1999).

Recently, we prepared a new class of polymeric micelles

(polymer –metal complex micelles) incorporating CDDP

through the polymer –metal complex formation between

CDDP and PEG-b-P(Asp) ( Nishiyama et al., 1999, 2001) or

PEG-block -(glutamic acid) [PEG-b-P(Glu)] copolymers (Fig.

4B) ( Nishiyama et al., 2003d). This micelle formation is based

on the ligand substitution reaction of the Pt(II) from chloride(leaving group) to carboxylate in the block copolymers (Fig.

4C). The CDDP-loaded micelles have a diameter of ca. 30 nm

with a narrow size distribution and showed a remarkable

stability in distilled water. In physiological saline (0.15 M

NaCl), however, the inverse ligand substitution reaction of Pt

(II) from the carboxylate to chloride ions in the medium occurs,

so that the micelles slowly release CDDP, accompanied by the

dissociation of the micellar structure with an induction period of

ca. 10 hr (Fig. 4D). When intravenously injected into tumor-

bearing mice, CDDP-loaded micelles showed > 60% of injected

dose in the plasma up to 8 hr and 13% of the plasma Pt level

even at 24-hr post-injection (65-folds higher AUC of Pt concentration–time curve than free CDDP), thereby resulting

in their significant accumulation in solid tumors (a 20-fold

higher concentration than free CDDP) ( Nishiyama et al.,

2003d). The accumulation and AUC ratios of the tumor to

normal tissues at 24-hr post-injection are summarized in Table

1. Free CDDP showed < 0.13 and 0.38 for the ratios of the tumor

to the kidney and liver, respectively, indicating its specificity to

the kidney and liver. In contrast, the CDDP-loaded micelles

showed the accumulation and AUC ratios higher than 1.0,

suggesting their selective accumulation in the tumor. Conse-

quently, the CDDP-loaded micelles exhibited in vivo antitumor

activity equivalent to or better than free CDDP depending on

the tumor models, while showing a significantly reducednephrotoxicity and neurotoxicity, which were confirmed by

histological and functional analyses ( Nishiyama et al., 2003d;

Uchino et al., 2005). Such reduced nephrotoxicity and

neurotoxicity of the CDDP-loaded micelles seemed to be

attributable to their significantly decreased maximum concen-

tration (C max) in the kidney and reduced cumulative concentra-

tion in the peripheral nerve, respectively, compared to free

CDDP. It should be noted that the treatment with the CDDP-

loaded micelles caused a transient and reversible hepatic

dysfunction, raising some caution to its long-term administra-

tion (Uchino et al., 2005). CDDP-loaded micelles are expected

to be a promising formulation of CDDP for clinical cancer chemotherapy.

1,2-Diaminocyclohexane platinum (II) (DACHPt) com-

plexes are a new class of platinous drugs showing a wide

spectrum of activity different from CDDP, and characterized by

a lower toxicity and no cross-resistance with CDDP in many

CDDP-resistant cancers (McKeage, 2005). Among them,

oxalate 1,2-diaminocyclohexane platinum(II) (oxaliplatin) has

been approved for the treatment of malignant tumors including

colorectal cancer (Chau & Cunningham, 2003). Recently, we

have prepared the DACHPt-loaded polymeric micelles with a

size of 40 nm through the polymer –metal complex formation

between DACHPt and PEG-b-P(Glu) using a method similar to

the CDDP-loaded micelles (C ab ral et al. , 2 00 5) . I n




physiological saline, the DACHPt-loaded micelles showed the

sustained release of the Pt(II) complexes with an induction

period of 15 hr. Also, the DACHPt-loaded micelles maintained

their micellar structure for a >5-fold prolonged period

compared to the CDDP-loaded micelles in spite of their

comparable release rate. These results may be related to the

hydrophobic nature of the DACHPt complexes in the micellar core. It was also demonstrated that the DACHPt-loaded

micelles show a prolonged circulation period (13% of injected

dose at 24-hr post-injection) after i.v. injection and thereby an

enhanced accumulation in the solid tumors (in a >10-fold

higher concentration than oxaliplatin) (Cabral et al., 2005).

In these systems, the ligand substitution of the leaving group

of the Pt(II) complexes may possibly affect their pharmacolog-

ical activities. In this regard, the patterns of 50% growth

inhibitory concentrations (IC50) against 39 different cancer cells

(fingerprints) were investigated (unpublished data). The

fingerprints of the CDDP-loaded micelles show similarity

with those of CDDP (r =0.906) and carboplatin (r =0.662). On

the other hand, the similarity was found between the DACHPt-

Table 1

Accumulation ratios and AUC ratios of the tumor (Lewis lung carcinoma cells)

to normal organs at 24-hr post-injection of free CDDP and the CDDP-loaded

micelle (CDDP/m)

Accumulation ratio a AUC ratio b

CDDP CDDP/m CDDP CDDP/m

Tumor/kidney 0.13 ±0.02 2.0 ±0.4 0.13 0.97

Tumor/liver 0.34 ± 0.07 1.6 ± 0.3 0.38 1.3

Tumor/spleen 4.0 ± 1.5 1.3 ± 0.1 2.4 1.5

Dose: 0.1 mg/mouse on CDDP basis (Copyright 2003 American Association for

Cancer Research).a

Mean±SE (n =4). b AUC was calculated based on the trapezoidal rule up to 24 hr.

Fig. 4. Chemical structures of CDDP (A) and PEG-b-P(Glu) copolymers (B), and schematic illustrations of CDDP-loaded micelles (C) and the hypothesized behavior

of the micelles in physiological saline (D). The CDDP-loaded micelles are spontaneously formed via the ligand exchange reaction of Pt(II) from the chloride to the

carboxylates in the copolymers in distilled water (C), and the micelles dissociate accompanied with the sustained release of CDDP via an inverse ligand exchange

reaction of Pt(II) from the carboxylates in the copolymer to the chloride ions in the surroundings in physiological saline (D) (Copyright 2003 American Association for

Cancer Research).




loaded micelles and oxaliplatin (r =0.785). Thus, the CDDP-

and DACHPt-loaded micelles maintained the pharmacological

properties (e.g., no cross-resistance) of the parent drugs.

3.4. Smart polymeric micelles for site-specific drug delivery

There has recently been a strong impetus to the development of polymeric micelles with smart functions, such as targetability

to specific tissues ( Nagasaki et al., 2001; Lee et al., 2003;

Torchilin et al., 2003; Bae et al., 2005a, 2005b) and chemical

(Bae et al., 2003; Lee et al., 2003; Bae et al., 2005b ) or physical

(Kohori et al., 1999) stimuli-sensitivity (Fig. 1). Smart

polymeric micelles are aimed to increase the selectivity and

efficiency of drug delivery to the target cells, leading to a better

therapeutic efficacy as well as reduced side effects. In this

section, the rationale to design such smart micelles is briefly

reviewed.

A minimal leakage of the loaded drugs from drug carriers

during the blood circulation should ensure the safety of chemotherapy, facilitating the development of nanocarriers to

release the loaded drugs selectively inside the cells. Indeed,

several intracellular signals, such as low pH (Bae et al., 2003;

Lee et al., 2003; Bae et al., 2005a, 2005b), glutathion (Miyata et

al., 2004) and specific enzymes (Katayama et al., 2001), have

been so far used for designing the environmentally sensitive

polymeric nanocarriers. Recently, we have developed pH-

sensitive polymeric micelles, in which Dox is attached to the

side chain of the core-forming segment of the block copolymers

via an acid-labile hydrazone bond (Bae et al., 2003, 2005a,

2005b). The pH-sensitive polymeric micelles showed a

significant drug release under endosomal/lysosomal low pH

conditions (5.0–5.5), while exhibiting no appreciable releaseunder physiological pH conditions. Such a high selectivity of

the drug release is attributable to a >100-fold difference in the

proton concentration between the endosomal/lysosomal com-

partment and the extracellular medium. A biodistribution study

revealed that the pH-sensitive polymeric micelles show a

prolonged blood circulation due to a minimal leakage of free

drug, resulting in the highly selective accumulation in solid

tumors. As a result, the pH-sensitive polymeric micelles

achieved a significantly higher antitumor activity in C-26-

bearing mice over a broader range of injection doses than free

Dox (Bae et al., 2005a).

Highly selective drug delivery can be achieved by theabove-described intracellular environmentally selective drug

release coupled with a specific drug delivery to the target

tissue. In particular, specific drug delivery to cancer cells

requires the following processes: prolonged blood circula-

tion, extravasation and local retention in tumor tissues. As

already mentioned, particle size and surface properties of

drug carriers might predominantly affect all these processes.

On the other hand, the modification of drug carriers with

specific ligands to cancer cells (i.e., active targeting) is

assumed to contribute to their local retention in tumor

tissue, and also promote their internalization through the

receptor-mediated endocytosis. Indeed, it was demonstrated

that transferrin-conjugated long-circulating liposomes show a

remarkably prolonged residence in tumor tissues compared

with the non-targeted long-circulating liposomes, in spite of

their similar profiles of blood clearance and tumor

accumulation (Maruyama et al., 1999). Thus, tumor-targeted

therapy using the targetable nanocarriers is feasible,

especially against readily accessible tissues such as lung

cancers (Maruyama et al., 1999). Recently, we modified theaforementioned pH-sensitive polymeric micelles incorporat-

ing Dox with a folate molecule to construct polymeric

micelles with tumor-targetability and intracellular pH-sensi-

tivity (Bae et al., 2005b). It is known that a large number of

cancer cells overexpress folate-binding proteins (FBP)

(Weitman et al., 1992). The folate-conjugated micelles

specifically bound to FBP immobilized on a dextran-coated

sensor chip in the surface plasmon resonance (SPR)

measurement, and were more efficiently taken up by the

FBP-overexpressing human pharyngeal cancer KB cells than

the non-targeted micelles. In the cytotoxic activity assay

against KB cells, folate-conjugated micelles showed acomparable cytotoxicity to free Dox after a 24-hr exposure

time, whereas the non-targeted micelles had almost a 10-fold

lower cytotoxicity than free Dox (Bae et al., 2005b). It is

unprecedented that the folate-conjugated micelles achieved a

high cytotoxicity as free Dox despite their different

internalization pathways. This result indicates that the use

of the folate-conjugated micelles may lower the effective

doses over free Dox, improving the safety of the clinical

chemotherapy.

In addition to the selective binding to cancer cells, actively

targeted polymeric micelles need to penetrate into the avascular

tumor tissues for eradication of the solid tumors. However, the

tumor-infiltrating ability of polymeric nanocarriers seems to becontroversial. In this regard, we have recently evaluated the

tissue penetration of polymeric micelles charged with Dox via

an acid-labile bond and subsequent drug release inside the

multicellular tumor spheroids with a diameter of 200 μm (Bae et

al., 2005a), because a tumor spheroid is known as an

appropriate in vitro experimental model of avascular tumor

tissues (Sutherland, 1988; Hamilton, 1998). It is noted that the

furthest distance between adjacent capillaries in an avascular

solid tumor is 200 μm or less (Konerding et al., 2001). As a

result, the Dox-loaded micelles were found to access every cell

inside the spheroids within a 3-hr incubation, followed by an

appreciable drug release inside the cell (Bae et al., 2005a).These results suggest a high tumor-infiltrating ability and

intracellular pH-triggered drug release of the polymeric

micelles. In contrast, PEG-liposomes incorporating Dox

(Doxil) showed a moderate fluorescence limited to the

periphery of the spheroids even after a 24-hr incubation,

suggesting limited accessibility to the inside of the spheroids

and difficulty in drug release (Tsukioka et al., 2002). The larger

size of Doxil (100–150 nm) compared to the polymeric micelles

(∼65 nm) may account for such a difference in their tumor-

infiltrating ability. Such a high tissue penetrating property of the

polymeric micelles may be promising for the active targeting of

solid tumors as well as offering the potential treatment of

avascular sites (hypoxic regions) of tumor tissues.




3.5. Effects of polymeric

nanocarriers on intracellular drug action

Polymeric nanocarriers were originally designed to deliver

drugs to the target tissue in order to improve the therapeutic

effect and minimize the side effects. In addition to such control

of the biodistribution, the effects of drug carriers on theintracellular drug action have recently been attracting another

focus of attention. Drug carriers might change the subcellular

localization of drugs due to their characteristic internalization

mechanisms mainly through the endocytic pathway, and also

control the time-dependent intracellular concentration of the

active agents through a controlled drug release. These effects

may affect the pharmacological activity of the loaded drugs.

Also, there is another possibility that polymeric carriers or the

constituent polymers themselves may interact with cellular

components, altering the cellular response to the active agents.

However, such effects of the polymeric nanocarriers are still

unclear.It was reported that the use of polymeric carriers could

overcome multidrug resistance of cancer cells, which has been

hypothetically explained by circumvention of the drug efflux

pumps (e.g., P-glycoprotein) overexpressed on the plasma

membrane through the endocytic pathway (Miyamoto &

Maeda, 1990; Omelyanenko et al., 1998). However, other

effects of the polymeric carrier remained to be studied, until

Minko et al. suggested that polymer –drug conjugates might

induce different mechanisms of cell death from the parent drugs

in vitro (Minko et al., 1999) and in vivo (Minko et al., 2000).

They found that the treatment with the N -(2-hydroxypropyl)

methacrylamide copolymer (PHPMA)–Dox conjugate

(PHPMA–Dox) downregulated genes responsible for the drugdetoxification and DNA repair ( HSP-70, GST-p, BUDP , Topo-

II α ,β , and TK-1), whereas free Dox treatment upregulated them.

Also, PHPMA–Dox more significantly upregulated the p53,

Apaf-1 and Caspase-9 genes than free Dox, and downregulated

the Bcl-2 gene which was upregulated by free Dox, suggesting

the elevated activation of the apoptotic signals by the PHPMA–

Dox treatment. Furthermore, PHPMA–Dox downregulated the

MDR-1 and MRP genes encoding the ATP-driven drug efflux

pumps, whereas free Dox upregulated them. Consequently,

PHPMA–Dox overcame multidrug resistance and induced

apoptosis as well as necrosis more efficiently than free Dox in

both the Dox-sensitive and -resistant cancer cells. Suchdifferences in the drug action between free and polymeric

carrier-loaded drugs have also been reported for other drugs

( Nishiyama et al., 2003c) and carrier systems (Minko et al.,

2005).

Recently, we also found that the CDDP-loaded polymeric

micelles (see Section 3.3) could regulate different genes from

free CDDP in human non-small cell lung cancer (NSCLC) cells.

For this purpose, we used the cDNA gene expression array

techniques (807 genes) ( Nishiyama et al., 2003b). The

hierarchical clustering analysis indicated that the total gene

expression profile was time-dependently approximated between

free CDDP and CDDP-loaded micelles (Fig. 5), which appears

to be consistent with the sustained CDDP release from the

micelles. Nevertheless, 50 genes were found to be differentlyregulated between free and micellar CDDP treatments (Fig. 5),

which included a number of important genes for the action of

CDDP, such as Chk1, PLC-delta and MDM2. It was suggested

that the treatment with the CDDP-loaded micelles activated

these genes towards anti-apoptosis, whereas free CDDP

activated them towards apoptosis. Interestingly, the CDDP-

loaded micelles downregulated the gene expression of the

integrin and matrix metalloprotease (MMP) families, which

could be possible targets related to angiogenesis and metastasis

for cancer treatment, whereas free CDDP upregulated them.

This result may offer additional therapeutic effects of the

CDDP-loaded micelles.As already mentioned, the interaction of the polymeric

carriers or the constituent polymers with cellular components

may lead to the altered cellular response to the active agents.

Pluronic® (poloxamers), an A-B-A type amphiphilic triblock

copolymer consisting of ethylene oxide (A) and propylene

oxide (B) segments, assembles into micelles in an aqueous

medium, and the Pluronic micelles have been used as the

carriers of hydrophobic drugs. Interestingly, Pluronic unimers

might interact with cell membranes and membrane proteins,

such as transporters, affecting their biological functions. It has

been suggested that Pluronic unimers could inhibit the P-

glycoprotein and other tansporters, leading to the sensitization

of the multidrug resistant cancer cells (Batrakova et al., 2001;

Fig. 5. Hierarchical cluster analysis of the expression of 807 genes related to cell

cycle regulation, oncogenes, angiogenesis, growth factors, cytokines, apoptosis-

related genes, and DNA transcription factors etc., on the gene expression array.

PC-14 cells were treated with free CDDP or CDDP-loaded micelle (CDDP/m) at

90% growth inhibitory concentrations (IC90) for 6 or 12 hr, followed by isolation

of total RNA (Copyright 2003 American Chemical Society).




Kabanov et al., 2005). Also, they recently found that Pluronic

block copolymers could increase the transgene expression after

the transfection using viral or nonviral vectors as well as

injection into muscles or tumors (Kabanov et al., 2005). The

Pluronic-mediated enhancement of the gene expression in the

muscles was comparable to the expression achieved by

electroporation (Lemieux et al., 2000). Pluronic also enhancedthe gene expression even toward stably transfected cells, and the

gene expression levels significantly depend on the promoter

type of plasmid (Kabanov et al., 2005). Therefore, it is

hypothesized that the Pluronic treatment may affect the

transcriptinal control of the transgene expression; however,

the mechanisms have not been clarified yet. Thus, the Pluronic

block copolymers are assumed to act as a biological response

modifier.

A significant focus has been recently placed on the control of

the subcellular localization of the polymeric nanocarriers.

Maysinger et al. recently reported that the micelles from the

PEG-b-poly(ε-caprolactone) copolymers are localized not onlyin the lysosome, but also in the mitochondrion, Golgi apparatus

and endoplasmic reticulum (Savic et al. , 2003). They

hypothesized that the micelles may dissociate into block

copolymers in the lysosome and perturb the lysosomal

membrane in order to relocalize the micelles. On the other

hand, modification of the polymeric carriers with peptides and

antibodies can actively control their subcellular trafficking

(Sheff, 2004). In particular, cell membrane-penetrating pep-

tides, such as the Tat peptide, are of significant interest because

the modified nanoparticles have shown to undergo their direct

and energy-independent transduction into the cytoplasm (Lewin

et al., 2000; Liu et al., 2001). Coupled with these technologies

to control the subcellular trafficking of the polymeric carriers, a better understanding of the intracellular action of the nanocar-

rier-loaded drugs will lead to the optimal design of the

nanocarriers.

4. Dendritic photosensitizer-loaded

polymeric micelles for photodynamic therapy

Photodynamic therapy (PDT) is a promising approach for the

treatment of malignant tumors and macular degradation

(Dougherty et al., 1998; Macdonald & Dougherty, 2001;

Renno & Miller, 2001). PDT involves the systemic adminis-

tration of photosensitizers (PS), followed by the localapplication of a laser with a specific wavelength to the diseased

sites. Upon photoirradiation, PS generate highly reactive singlet

oxygen (1O2), thereby inducing light-induced cytotoxicity

(photocytotoxicity). In PDT, the development of delivery

systems for PS has recently received much attention to improve

the selectivity and effectiveness of PDT as well as prevent the

side effects such as skin hypersensitivity. For example, the

polymer –PS conjugates (Tijerina et al., 2003), PEG-liposome

(Derycke & de White, 2004) and polymeric micelles (Le Garrec

et al., 2002) have been studied as a vehicle of PS. However, it is

generally difficult to effectively incorporate PS into drug

carriers, because they easily form aggregates through their

Π–Π stacking and hydrophobic interaction. Also, such an

aggregate formation of dendritic porphyrins (DP) is known to

significantly reduce the efficiency of the singlet oxygen

produ ction due to self-quenching of their excited state

(Grossweiner et al., 1982). Hence, both the efficiencies of the

PS delivery and the photochemical reactions of PS should be

taken into consideration in order to develop an effective

formulation for PDT.

Recently, we developed an ionic DP, in which the focal core

of the porphyrin is surrounded by the 3rd generation of poly

(benzyl ether) dendrons with peripheral ionic (carboxyl) groups(Fig. 6), as potential PS for PDT ( Nishiyama et al., 2003a). The

dendritic framework of the DP is assumed to sterically prevent

the interaction (i.e., self-quenching) of the center porphyrins,

ensuring the effective singlet oxygen production from DP even

at extremely high concentrations. Also, 32 carboxyl groups on

the periphery of DP allowed its stable incorporation into PIC

micelles through the electrostatic interaction with the positively

charged PEG-block -poly(L-lysine) (PEG-b-PLL) copolymers

(Fig. 7) (Stapert et al., 2000). Simple mixing of DP and PEG-b-

PLL resulted in the formation of narrowly distributed PIC

micelles with the diameter of ca. 60 nm. On the subject of the

photochemical reactions of DP, the DP-loaded micelles showedan oxygen consumption rate comparable to free DP in

phosphate buffered saline containing 10% fetal bovine serum

upon photoirradiation, although each micelle contains an

average of 38 DP molecules in the core (Jang et al., 2005).

No quenching of DP inside the micellar core is attributable to

the steric hindrance of the interaction between the dye

molecules by the dendritic framework. It is noteworthy that

serum proteins play the role as a singlet oxygen acceptor in the

oxygen consumption measurement. Hence, the singlet oxygen

molecule appears to reach the outside of the micelles to react

with the serum proteins, since proteins are immiscible in the

PEG layers. Interestingly, the DP-loaded micelles showed a

280-fold increase in the photocytotoxicity to Lewis lung

Fig. 6. Chemical structure of ionic dendritic porphyrin (DP) (X=COO− Na+).




carcinoma cells compared to free DP (Jang et al., 2005). It

appears that the unprecedented production of the singlet oxygenfrom the micellar core may significantly increase the oxidizing

levels of the subcellular molecules in the cell, leading to the

enhanced photocytotoxicity.

Exudative age-related macular degeneration (AMD), which

is characterized by CNV, is a leading cause of visual loss in

developed countries (Renno & Miller, 2001). PDT is known to

be effective in the treatment of AMD as Visudyne®, a liposomal

formulation of verteporfin that has recently been approved for

clinical use (TAP and VIP Study Group, 2002). However, most

patients require repeated treatments; therefore, the improvement

of the PDT efficacy is strongly demanded. Recently, we

attempted to treat the experimental CNV in rats with PDT using

the DP-loaded micelles. Microscopic observation of the frozen

tissues revealed that the DP-loaded micelles specifically

accumulated in the CNV sites (Fig. 8A) (Ideta et al., 2005).Probably, the CNV lesions may have the feature of vascular

hyperpermeability similar to solid tumors. Consequently, the

application of the laser 0.25 or 4 hr after injection resulted in a

60–78% occlusion of the CNV lesions (Fig. 8B, Table 2) (Ideta

et al., 2005). Importantly, an approximate 80% occlusion of the

CNV was maintained 7 days after the treatment, indicating the

effectiveness of PDT using the DP-loaded micelles (Table 2).

Immunohistochemical analysis and TEM observation of the

tissue section treated with PDT demonstrated that the CNV EC

are destroyed, and vessels in the CNV lesions become cell-free

collagen tubes or occluded by erythrocytes. Furthermore, the

PDT using the DP-loaded micelles resulted in no skin damage

when the rats were exposed to broadband visible light

Fig. 7. Formation of DP-loaded micelle through the electrostatic interaction between DP and PEG- b-PLL.

Fig. 8. (A) Accumulation of the DP-loaded micelles in CNV lesion in rat AMD model at 4-hr post-injection. The fluorescence of DP was detected in the factor VIII-

positive endothelial cells, and was still evident at 24-hr post-injection. (B) Occlusion of CNV in the rat eye by PDT using the DP-loaded micelles. The

hypofluorescence of i.p. injected fluorescein indicates successful occlusion of CNV (fluorescein angiography) (Copyright 2005 American Chemical Society).




simulating sunlight (Ideta et al., 2005). In contrast, the PDT

using Photofrin, a clinically used PS formulation resulted in

severe photodamage to the skin. Thus, the DP-loaded micelles

are expected to be an innovative PS formulation for the

enhanced PDT.

DP is applicable only to ophthalmic applications because of

its relatively short excitation wavelengths (430 and 559 nm).

For the PDT of solid tumors, dendritic PS with longer excitationwavelengths need to be developed. In this regard, we have

prepared PIC micelles encapsulating dendritic phthalocyanine

(DPc), which can be excited at 680 nm, and their applications

for the PDT of solid tumors are now ongoing. These results will

be reported elsewhere in the near future.

5. Nanodevices for gene therapy

5.1. Polyion complex (PIC) micelles for plasmid DNA delivery

Gene therapy is a promising approach for the treatment of

genetic and intractable diseases, and its success relies on the

capacities of gene vectors (Verma & Somia, 1997; Pack et al.,2005). Compared with viral vectors, non-viral gene carriers

have many advantages, such as safety for clinical use, simplicity

of preparation, and easy large-scale production. In this regard,

cationic polymers have been studied as non-viral gene carriers

due to their ability to package the negatively-charged plasmid

DNA (pDNA) into a small particle (< 200 nm) for the protection

of DNA from enzymatic and hydrolytic degradation as well as

the effective cellular uptake through the endocytosis. However,

the pDNA/cationic polymer complexes (called “ polyplexes”)

may not be useful for in vivo gene delivery due to their cationic

property, which might lead to an uncontrollable biodistribution

in the body (Ward et al., 2001) and may cause fatal toxicityassociated with the occlusion of the lung capillaries via

erythrocyte aggregation (Boeckle et al., 2004). A promising

way to improve the biocompatibility of the polyplexes is the use

of PEG-b-polycation copolymers, which electrostatically inter-

act with pDNA to form the PIC micelles (Fig. 1). For instance, a

simple mixing of pDNA and PEG-b-PLL at the Lys/nucleotide

unit ratio of 2 resulted in the spontaneous formation of the

pDNA-loaded micelles characterized by a small particle size

(ca. 100 nm), low absolute ζ-potential value and excellent

colloidal stability (Katayose & Kataoka, 1997; Itaka et al.,

2003). The pDNA/PEG-b-PLL micelles maintained their

structures (condensed state of pDNA in the PIC core) and

gene transferring ability after incubation in serum-containing

media, which might be attributable to a unique core–shell

structure of the PIC micelles that reduces their interaction with

serum proteins (Itaka et al., 2002, 2003). When the pDNA/PEG-

b-PLL micelles were intravenously injected, an intact pDNA

was observed in the blood circulation even at 3-hr post

injection, which is in marked contrast with the injection of

naked pDNA being eliminated from the circulation within 5 min(Harada-Shiba et al., 2002). Thus, it appears that the PIC

micelles are suitable for in vivo gene delivery.

The PIC micelles may need to be further stabilized under

harsh in vivo conditions, where abundant negatively charged

macromolecules such as albumin exist and may destabilize the

PIC by the counter polyelectrolyte reaction. To further

improve the stability of the PIC micelles, therefore, the PLL

segment of PEG-b-PLL was modified with the thiol group,

thereby crosslinking the PIC core through the formation of

disulfide bonds (Kakizawa et al., 1999; Miyata et al., 2004,

2005). The disulfide bonds are assumed to be selectively

cleavable in the cytoplasm, because the glutathione concen-tration in the cytoplasm is 50–1000 times higher than that in

extracellular media. Indeed, the crosslinked micelles showed

an efficient pDNA release responding to the reductive

conditions mimicking the intracellular environment, thereby

inducing better transfection to the cultured cells than the non-

crosslinked micelles. Such stabilization of the PIC micelles

through the disulfide crosslinking might be useful for the

systemic gene delivery, because crosslinked PIC micelles are

expected to be stable during blood circulation, but release

pDNA inside the targeted cells through the cleavage of the

disulfide bonds. Indeed, the intravenous injection of cross-

linked PIC micelles into mice resulted in a uniform gene

expression in the liver (Miyata et al., 2005).To achieve a site-specific gene delivery, polyplex micelles

might be modified with targetable ligands such as peptides ( Nah

et al., 2002) and antibodies (Vinogradov et al., 1999; Merdan et

al., 2003). Recently, we introduced a lactose moiety into the

distal end of PEG on the polyplex micelles for the hepatocyte-

specific gene delivery. The lactosylated polyplex micelles

showed an enhanced gene transfection to asialoglycoprotein

(ASGP) receptor-positive human hepatoma HepG2 cells

compared to the non-targeted polyplex micelles (Wakebayashi

et al., 2004). The addition of excess asialofetuin, a natural

ligand against the ASGP receptor, resulted in a significant

decrease in the gene transferring activity of the lactosylatedmicelles, suggesting the internalization of the lactosylated

micelles via receptor-mediated endocytosis. The targetable

polyplex micelles have a great potential for the site-specific

gene delivery via systemic administration.

Despite the aforementioned endeavors to deliver the

therapeutic genes to the target tissue, the improvement of the

transfection activity of the polyplex micelles is an important

issue to be addressed for their clinical applications. The

polyplex micelles are assumed to be taken up by the cell via

the endocytic pathway, ending in localization in the lysosome;

therefore, the endosomal escape of the polyplex micelles before

reaching the endosome might be a key to the enhancement of

the transfection efficiencies. In this regard, cationic polymers

Table 2

Occlusion efficiency of CNV after PDT using the DP-loaded micelle

Light

fluence a

(J/cm2)

Control PDT using DP-loaded

micelle

Day 1 Day 7 Day 1 Day 8

0 31.7 ± 13.4 15.0 ± 3.5 33.3 ± 0 25.0 ± 8.3

5 60.0 ± 10.0 81.7 ± 1.750 72.2 ± 15.5 77.8 ± 2.8

(Copyright 2005 American Chemical Society).a PDT laser was applied 4 hr after i.v. injection of the DP-loaded micelle.




with a comparatively low p K a , such as polyethylenimine (PEI),

are known to be highly transfectable, because they could buffer

the endosomal acidification as well as cause an increase in the

ion osmotic pressure in the endosome accompanied by the

protonation of amines, disrupting the endosomal membrane to

release its contents into the cytoplasm. Such effects of the

cationic polymers with a low p K a are called the “ proton spongeeffect ” (Behr, 1997), and have been an important basis for

designing the polyplexes as non-viral gene vectors. However,

the polyplexes from such low p K a polycations require excess

polycations to provide a high stability and efficient gene

transfection. Recently, Boeckle et al. (2004) demonstrated that

the PEI polyplexes contain free PEI, which substantially

contributes to the enhanced gene expression as well as the

cytotoxicity. Such polyplexes containing free polycations might

not be useful for the systemic gene delivery due to instability

and toxicity problems.

To integrate the proton sponge effect into the polyplex

micelle system, we recently developed an A-B-C type triblock copolymer consisting of PEG, poly[(3-morpholinopropyl)

aspartamide] (PMPA) as a low p K a polycation and the PLL

segment (PEG-b-PMPA-b-PLL) (Fukushima et al., 2005). In

the triblock copolymer, the PLL segment preferentially interacts

with the negatively charged DNA, allowing the formation of the

3-layered polyplex micelles where an inner core of the pDNA/

PLL polyplex is successively wrapped with an intermediate

layer of the low p K a PMPA segment, and an outer layer of the

biocompatible PEG segment, as illustrated in Fig. 9. The PEG-

b-PMPA-b-PLL polyplex micelles showed a >10-fold higher

transfection efficiency to human hepatoma HuH-7 cells than the

polyplexes from PEG-b-PLL or the mixture of PEG-b-PLL and

PEG-b-PMPA due to the high buffering capacity of the PMPAsegment remaining free in the intermediate layer of the 3-

layered micelles. It is noteworthy that such an enhancement of

the gene transfection of the polyplex micelles was achieved

under the conditions without free polymers. In the triblock

copolymers, a high buffering capacity and the stabilization of

the polyplexes, both of which are essential to nanocarriers for

systemic gene delivery, are assigned to separate cationic

polymers with different p K a in a single polymer strand,

accounting for the aforementioned enhancement of the

transfection activity of the polyplex micelles under the

conditions of minimal free polymers. Further studies on the in

vivo applications of the 3-layered polyplex micelles are now

ongoing in our laboratory.

5.2. PIC micelles for small interfering RNA delivery

siRNA are recognized as the most powerful tool for silencing

the gene expression in a sequence-specific manner (Elbashir et

al., 2001), and their therapeutic applications have received the

utmost interest in recent years. Nevertheless, the lack of

appropriate carrier systems for the in vivo siRNA delivery

remains a limitation for clinical applications (Pack et al., 2005).

Also, nanocarrier systems are required to improve the fragility,

impermeability to the cellular membranes and the undesirable

biodistribution of siRNA. Polymeric micelles might be a useful

candidate for siRNA nanocarriers.

A negatively charged siRNA can be incorporated into PIC

micelles through the electrostatic interaction with PEG-b- polycation block copolymers. In our recent study, a PEG-b-

polycation possessing a diamine structure with 2 distinct p K a ,

that is, primary and secondary amino groups, in the side chain

(PEG-b-DPT), was found to be remarkably effective for the

siRNA delivery (Itaka et al., 2004). This unique structure of

PEG-b-DPT may allow only the primary amino group to be

involved in the PIC formation, thereby maintaining a buffering

capacity of the secondary amino group for the proton sponge

effect. The PIC micelles of PEG-b-DPT/siRNA showed a

significant gene silencing toward endogenous genes (e.g.,

Lamin A/C) even after a 30-min pre-incubation in 50% serum.

These properties of the PIC micelles offer a promising

feasibility for in vivo siRNA delivery.Alternatively, PIC micelles were formed between the PEG-

b-siRNA conjugates and polycations. We have recently

conjugated siRNA with lactosylated PEG through an acid-

labile linkage of the β-thiopropiopnate to obtain Lac-PEG-b-

siRNA, followed by complexation with the PLL homopoly-

mers to form the lactosylated PIC micelles (Fig. 10) (Oishi et

al., 2005). Such PIC micelles of Lac-PEG-b-siRNA/PLL are

assumed to be internalized through the ASGP receptor-

mediated endocytosis and thereafter exert the siRNA activity

Fig. 9. Chemical structure of PEG-b-PMPA-b-PLL triblock copolymers and schematic illustration of the 3-layered polyplex micelles with spatially regulated structure(Copyright 2005 American Chemical Society).




triggered by the cleavage of the β-thiopropiopnate bond under intracellular low pH conditions. In the dual luciferase reporter

assay using ASGP receptor-positive human hepatoma HuH-7

cells, the PIC micelles of Lac-PEG-b-siRNA/PLL showed a

50% gene silencing at 1.3 nM siRNA, which was remarkably

lower than the IC50 of Lac-PEG-b-siRNA alone (91.4 nM) and

the PIC micelles from the PEG-b-antisense DNA conjugate

(7.6 μM). The addition of excess asialofetuin resulted in a

significant decrease in the siRNA activity of the lactosylated

PIC micelles while showing no effect on the activity of Lac-

PEG-b-siRNA alone, suggesting the importance of the lactose

ligand clustering on the PIC micelles to facilitate the ASGP

receptor-mediated endocytosis. Worth mentioning is that high

molecular-weight polycations can be used for the PEG-b-siRNA/polycation system, expecting that the critical micelle

concentration (c.m.c.) of the PIC micelles may be remarkably

lowered compared to the aforementioned PEG-b-polycation/

siRNA system. This property is a significant advantage of the

PEG-b-siRNA/polycation system, because the siRNA carriersare extremely diluted during blood circulation after the sys-

temic administration.

Calcium phosphate (CaP)/DNA coprecipitation is a well-

known method for the transfection into mammalian and plant

cells, and this method might be useful for the siRNA delivery. It

is known that CaP including hydroxyapatite is one of the most

widely used biomaterials in biomedical applications due to its

excellent biocompatibility. However, uncontrollable growth of

the CaP crystal within tens of seconds results in difficult

handling and reproducibility. Recently, we have successfully

prepared CaP nanoparticles stabilized by PEG-b-P(Asp) block

copolymers as a novel nanocarrier of siRNA (Kakizawa et al.,

2004a, 2004b). The CaP nanoparticles with a core–shellstructure are prepared by mixing Ca2+ and PO4

3- ions in the

presence of siRNA and PEG-b-P(Asp) with different concen-

trations. During the CaP nanoparticle formation, the PEG

segment could sterically prevent the overgrowth of the CaP

Fig. 11. Formation of PEG-coated CaP nanoparticle incorporating siRNA. The PEG segment could prevent the overgrowth of CaP crystal, while P(Asp) segment and

siRNA are being incorporated into the core of CaP nanoparticle. The CaP nanoparticle may release siRNA selectively in the cytoplasm due to 20,000-folds lower Ca

2+

ion concentration than that of the extracellular fluid.

Fig. 10. Formation of lactosylated PIC micelles incorporating siRNA through the electrostatic interaction between PEG- b-siRNA conjugates and polycations. The

acid-labile bond between PEG and siRNA might be cleaved selectively under the endosomal low pH conditions.




crystal while the P(Asp) segment and siRNA being incorporated

into the nanoparticles through the interaction between the CaP

and polyanions. This thermodynamic control of the CaP

nanoparticle formation allows easy handling and ensures

reproducibility. The size of the CaP nanoparticles can be

controlled in the range of 100–300 nm by changing PEG-b-P

(Asp) and phosphate anion concentrations (Fig. 11). The

loading efficiency of siRNA was maintained at >89% even

for the remarkably high concentration of PEG-b-P(Asp),

although the competitive binding to CaP between PEG-b-P(Asp) and siRNA might occur. Interestingly, the CaP nano-

particles were dissolved selectively in the medium containing

an intracellular concentration of Ca2+ ion (∼100 nM), which is

20,000-folds lower than the calcium ion concentration in the

extracellular fluid (2 mM) (Clapham, 1995), allowing the

release of siRNA in an intracellular condition-selective manner.

The dual luciferase reporter assay revealed that the CaP

nanoparticles incorporating siRNA showed appreciable gene

silencing in a sequence-specific manner. Thus, the core–shell

type CaP nanoparticles are expected to be biocompatible

nanocarriers for siRNA delivery.

5.3. Novel gene carriers enveloped in

dendritic photosensitizer for light-induced gene transfer

The temporal and spatial control of the transgene expression

in the body is required to ensure the safety and effectiveness of

nonviral gene therapy; however, the existing vectors including

Fig. 12. Preparation of the pDNA/CP4/DPc ternary complexes and their

hypothetical mechanisms in the light-induced gene transfection ( Nishiyama et

al., 2005a).

Fig. 13. Transfection to the conjunctival tissue to rat eyes. (A) Scheme for in vivo transfection. Rats were given subconjunctival injection (colored in light blue) of the

ternary complexes, followed by the laser irradiation to a part of the conjunctiva (red circle) at 2 h post-injection. ( B, C) Fluorescent images of the reporter gene

expression in the rat eye at 2 days after the PCI-mediated transfection ( Nishiyama et al., 2005a). (For interpretation of the references to colour in this figure legend, thereader is referred to the web version of this article.)




viral and non-viral ones might lack the ability to control the

gene expression. The use of external stimuli for the enhance-

ment of the transgene expression may be a promising approach

to the site-directed transfection in vivo. In this regard, a new

technology called “ photochemical internalization (PCI)” has

recently emerged, in which the endosomal escape of the

polyplexes is induced by co-incubated PS which photodamagethe endosomal membrane, allowing gene transfection in a light-

inducible manner (Berg et al., 1999; Høgset et al., 2000;

Prasmickaite et al., 2001; Høgset et al., 2002, 2004). This

strategy is quite smart; however, the emergence of significant

photocytotoxicity was also found, possibly limiting its further

applications for in vivo use.

To solve this problem, we assumed that the control of the

intracellular localization of PS should be of primary importance,

and the photosensitizing property is preferably integrated into

the gene delivery system. Recently, we developed novel ternary

complexes, in which the pDNA/polycation polyplex is envel-

oped with the anionic DPc for effective PCI-mediatedtransfection (Fig. 12) ( Nishiyama et al., 2005a). In this study, a

disulfide-linked cationic peptide containing a nuclear localiza-

tion signal (NLS), quadruplicated cationic peptide (CP4), was

used as the polycations for the pDNA condensation. It was

demonstrated that the pDNA/CP4/DPc ternary complexes might

undergo the following processes during the light-induced

transfection (Fig. 12): (i) cellular uptake of the ternary

complexes via endocytosis, (ii) dissociation of DPc from the

complexes in acidic vesicles due to the protonation of the

carboxyl groups on the dendrimer periphery and increased

interaction of DPc with the endosomal membrane, and (iii)

endosomal escape of the pDNA/CP4 complexes to the cytoplasm

upon photoirradiation. Also, the pDNA/CP4 complex may bedelivered to the nuclei due to the NLS function (Rudolph et al.,

2003). As a result, the ternary complexes achieved more than a

100-fold photochemical enhancement of the transgene expres-

sion in vitro with reduced photocytotoxicity. The subconjuncti-

val injection of the ternary complexes in rat eyes followed by the

laser irradiation resulted in an appreciable gene expression (a

variant of yellow fluorescent protein) only at the laser-irradiated

site (Fig. 13). These results are the first success of the PCI-

mediated gene delivery in vivo. We are now elaborating the

systemically injectable gene carriers with the light-inducible

gene transferring ability. These light-responsive gene carriers are

expected to be useful for the site-directed transfection in vivo.

6. Future prospects

This paper reviewed recent progress in research on

polymeric micelles as nanocarriers for drug and gene delivery.

Polymeric micelles encapsulate various drugs including hydro-

phobic compounds, metal complexes, gene and siRNA, and

their unique core–shell architecture with a diameter of several

tens of nanometers might allow prolonged blood circulation and

preferential accumulation in solid tumors. Importantly, the

critical features of polymeric micelles as drug carriers can be

modulated by engineering the constituent block copolymers.

Unlike PEG-liposomes, polymeric micelles might show a

tumor-infiltrating ability as well as controlled drug release,

which is likely to be essential for the eradication of a tumor

mass. Interestingly, it was found that polymeric micelles

accumulated not only in solid tumors but also in the balloon-

injured site in rat carotid arteries or the CNV site in rat eyes,

offering their potential utility for the targeted therapy of the

cardiovascular or ophthalmic diseases, respectively.The development of polymeric micelles with smart functions

such as the environment-sensitivity and specific tissue-target-

ability may enhance the activity of potent bioactive compounds,

facilitating their clinical applications. Also, polymeric micelles

responsive to external stimuli, such as light, might exert the

activity of the loaded compounds in a site-directed manner,

ensuring the effectiveness and safety of the nanocarrier-

mediated targeting therapy. Thus, polymeric micelle-based

nanocarriers will continue to hold a promise for the delivery of

drugs and genes.

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