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7/23/2019 Current State Achievements and Future Prospects of Polymeric Micelles as Nanocarriers for Drug and Gene Deliver…
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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
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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.
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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
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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
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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).
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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
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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).
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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.
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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).
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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+).
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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).
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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.
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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).
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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.
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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.)
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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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