In situ Gelation of Supramolecular Hydrogel for Anti-Tumor Drug Delivery

Full Paper

In situ Gelation of Supramolecular Hydrogelfor Anti-Tumor Drug Delivery

Bin He,* Jing Zeng, Yu Nie, Li Ji, Rui Wang, Yuan Li, Yao Wu, Li Li,Gang Wang, Xianglin Luo, Zhirong Zhang, Zhongwei Gu*

A supramolecular injectable hydrogel was fabricated. The hydrogel was in situ gelated by thehost-guest interaction between a-cyclodextrins (a-CDs) and methylated poly(ethylene glycol)grafted poly(a,b-malic acid) (mPEG-g-PMA). The hydrogel was characterized by 1NMR, XRD,DSC, TGA and SEM. The results showed that the polyrotaxanes of a-CDs/mPEG-g-PMA acted asphysical crosslink sites in the hydrogel. Anti-tumor drug doxorubicin hydrochloride (DOX)wasloaded in the hydrogel. The release and anti-tumoreffect were studied in vitro. The burst release ofDOX was restrained obviously. The sustaining releasetime lasted more than 3d and the cell viabilitydecreased greatly. This hydrogel is a promising inject-able hydrogel for minimally invasive therapeutic drugdelivery.

Introduction

Hydrogels are widely used in tissue engineering and drug

delivery systems for their excellent biocompatibility and

large amount of water content.[1,2] Injectable hydrogels are

attracting much interest from biomaterials scientists.[3–6]

The in situ gelation, which could be triggered by

temperature, pH, redox and so on, is a main approach to

prepare injectable hydrogels.[7–9] Recently, host-guest

interactions between cyclodextrins and polymeric chains

were reported for injectable hydrogel fabrication.[10]

B. He, Z. Gu, J. Zeng, Y. Nie, L. Ji, Y. Li, Y. Wu, L. Li, G. WangNational Engineering Research Center for Biomaterials, SichuanUniversity, Chengdu, 610064, ChinaFax: 0086 28 8541 0653; E-mail: [email protected];[email protected]. Wang, X. LuoSchool of Polymer Science and Engineering, Sichuan University,Chengdu, 610065, ChinaZ. ZhangWest School of Pharmacy, Sichuan University, Chengdu, 610041,China

Macromol. Biosci. 2009, 9, 1169–1175

� 2009 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

Cyclodextrins (CDs) are cyclic oligosaccharides composed

of 6 (a), 7 (b) or 8 (g) D-glucose units linked by 1,4-a-

glucosidic bonds. The average cavity diameters of a-, b- and

g-CD are 4.5 A, 7.0 A and 8.5 A. The nature of the CD cavities

facilitates the ability of CDs (hosts) to form polyrotaxanes

with polymers (guests).[11,12] With the self-assembling of

polyrotaxanes, supramolecular injectable hydrogels could

be fabricated. Poly(ethylene glycol) (PEG) and its copoly-

mers are the main backbones of supramolecular injectable

hydrogels. Injectable hydrogels composed of cyclodextrins

and PEG based tri-block linear copolymers, such as

Pluronics, PCL-PEG-PCL and PEG-PHB-PEG, have been

reported.[13–17] Other grafted, star-shaped and brushes like

PEG copolymers have also been designed, synthesized and

developed into injectable hydrogels.[18–21] The preferable

biomedical application of the hydrogels is drug delivery. Li

loaded fluorescein isothiocyanate labeled dextran (dextran-

FITC) as a model macromolecular drug into supramolecular

injectable hydrogels to investigate the drug release

behavior. It was found that both the host-guest and

hydrophobic interactions contributed to the sustaining

release of dextran-FITC, and the release rate was closely

related to the amount of a-CDs.[15,17,22] Chen introduced

DOI: 10.1002/mabi.200900225 1169

B. He et al.

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protonizable segments to endow dual temperature and pH

responsive functions to supramolecular injectable hydrogel

drug delivery systems. The results showed that the gelation

was influenced by concentration, pH, composition and

uniformity.[23] So far, drug release studies are limited to

macromolecular model drugs. There are rare further studies

on real drug release and therapeutic effects.

Biodegradability is a necessary property for drug delivery

systems. In previous studies, the biodegradable segments in

supramolecular injectable hydrogels were hydrophobic

polyesters such as PCL and PHB. Poly(malic acid) (PMA) is a

water soluble polyester with functional carboxyl pendant

groups and good biodegradability. The UV crosslinked

poly(a,b-malic acid) based biodegradable hydrogel has been

studied,[24] but the hydrogel was neither injectable nor

supramolecular in structure.

In this paper, a PMA based supramolecular injectable was

prepared. mPEG with different molecular weights was

grafted onto PMA. a-CDs were threaded onto mPEG chains

to self-assemble supramolecular injectable hydrogels. The

hydrogel structure was characterized by 1H NMR, XRD, DSC,

TGA and SEM. The anti-tumor drug doxorubicin was loaded

in the hydrogel. The toxicity of the hydrogel was

investigated, and the drug release behavior, and the anti-

tumor effect was studied in vitro.

Experimental Part

Materials

L-malic acid, N-N0-dicyclohexylarbodiimide (DCC) and a-cyclodex-

trin were purchased from Aldrich and used as received. Doxor-

ubicin hydrochloride (DOX) was purchased from Haizheng

Pharmacy (China). Methylated poly(ethylene glycol)s

(Mn ¼ 750 g �mol�1 and 2 000 g �mol�1) were purchased from

Aldrich and vacuum-dried before used. Tetrahydrofuran (THF),

diethyl ether, acetone and dimethylsulfone (DMSO) were pur-

chased from Sinopharm chemical reagent company. THF and

diethyl ether were dried by refluxing over sodium. DMSO-d6 was

purified by vacuum distillation before used. Acetone was used as

received. The purity of all the chemicals was AR.

Measurements

1H NMR spectra were recorded on a Bruker DMX-300 spectrometer,

working at 300.130 MHz. GPC was performed on a Waters Breeze.

The samples were measured at 35 8C with THF as the eluent at a

flow rate of 1.0 mL �min�1. X-ray diffractometry (XRD) patterns

were obtained at room temperature on a Rigaku D/max-2500 X-ray

diffractometer with a Cu Ka (l¼0.154 nm) radiation source. The

supplied voltage and current were set to 50 kV and 100 mA,

respectively. The samples were mounted on a sample holder and

scanned with a step size of 0.028 from 2 to 508. Differential scanning

calorimetric (DSC) measurements were performed on a TA System

Q2000 under nitrogen at a flow rate of 50 mL �min�1. Each sample

Macromol. Biosci. 2009, 9, 1169–1175


was heated from�80 to 100 8C at a heating rate of 10 8C �min�1 and

scanned twice to erase the thermal history. Thermal gravity

analysis (TGA) measurements were carried out on a TA System

Q500. The samples were heated from room temperature to 500 8C at

a heating rate of 10 8C �min�1 in a dynamic nitrogen atmosphere.

The morphologies of hydrogels were observed with scanning

electron microscopy (SEM) (Hitachi-S4800, Japan) at 10 kV. The test

for drug release concentration was carried out on UV Spectrometer

(Pekin Elmer, Lambda 650) at 490 nm. The phase contrast images

were taken using a Zeiss (Zeiss Axiovert 200M) inverted micro-

scope. All the hydrogel samples were freeze-dried before measure-

ment.

Synthesis of PMA

The synthesis of PMA was according to the method described in

ref.[24]. In brief, 50 g of L-malic acid was added to a 150 mL round-

bottomed flask with a magnetic stirrer. The polycondensation was

carried out under 0.1 mmHg in a vacuum at 110 8C for 72 h. The

white products were dissolved in anhydrous THF and precipitated

in a large amount of anhydrous diethyl ether. The diethyl ether was

removed and the remaining white precipitate was vacuum-dried at

room temperature for 24 h.

Preparation of mPEG-g-PMA

A typical procedure was as follows. Prescribed amounts of mPEG

and PMA were added to a 150 mL round-bottomed flask with

anhydrous THF. DCC (the mole ratio of DCC to mPEG was 2:1) was

dissolved in anhydrous THF and added to the mixture dropwise.

The mixture was magnetically stirred at 0 8C for 24 h. White

precipitate appeared in the mixture and the mixture was filtrated.

The filtrate was condensed and precipitated in distilled water. The

mixture was centrifuged and the solution was dried by lyophiliza-

tion. The obtained white powder was vacuum-dried at room

temperature for 48 h.

Hydrogel Formation and Drug Loading

Prescribed amounts of a-CDs, mPEG-g-PMA and DOX were

dissolved in distilled water. The solution was sonicated for 30 s

and kept at room temperature for 30 min. The hydrogel loaded with

drugs was formed.

Drug Release

The release experiment in vitro was carried out according to a

method reported previously.[25] 2 mL of hydrogel loaded with DOX

was put into dialysis tubes (Fisher Scientific, MWCO 2000 Dalton,

USA). Each dialysis tube was immersed into 50 mL of phosphate

buffer solution (PBS, 0.01 M, pH 7.4) and the media was constantly

stirred with 120 rpm at 37�0.5 8C. At specific time intervals, 0.5 mL

of medium was taken out and replaced with PBS. Five replicated

samples were recorded at each time point.

DOI: 10.1002/mabi.200900225

In situ Gelation of Supramolecular Hydrogel . . .

Table 1. The compositions of hydrogels.

Entry mPEG

Mw

Graft

Degree

Composition

% wt.-%

Copolymer a-CD DOX H2O

1 750 20 4.4 8.7 0.017 86.9

2 750 20 4.0 16.0 0.016 80.0

3 750 20 4.3 8.7 0.034 87.0

Cell Culture

The glioma cancer cells (U87MG) from American Type Culture

Collection (ATCC) were cultured in 50 mL cell culture flasks with

Dulbecco’s Modified Eagles Medium (DMEM, Gibco) buffered with

N-(2-hydroxyethyl)piperazine-N0-2-ethanesulfonic acid (HEPES),

supplemented with 15% calf serum (Gibco) and 100 U � cm�3

each of penicillin and streptomycin. The cell density was

1�105 cells �mL�1 and the cell culture was maintained in a gas

jacket incubator equilibrated with 5% CO2 at 37 8C. After 24 h

incubation, the blank hydrogels and hydrogels loaded with DOX

were co-cultured with the cancer cells. The cell culture medium was

changed every 2 d.
4 2 000 20 8.0 16.0 0.015 78.0
Cell Viability

The viability of U87MG cells was determined by 3-(4,5)-

dimethylthiahiazo(-z-y1)-3,5-di-phenytetrazoliumromide (MTT,

Sigma) assay. The MTT solution was 5 mg �mL�1 in PBS. The

solution was sterilized by Millipore filtration and kept in the dark.

100mL of MTT solution were added to each well of the cell culture

plate. After incubating at 37 8C for 4 h, 1 mL of DMSO-d6) was added

to dissolve the formazan crystals. The optical density (OD) of the

formazan solution was read on a microplate reader (BIO-RAD,

model 550, USA) at 570 nm. The cell viability was calculated as:

Macrom

� 2009

cell viability ¼ OD=OD0 � 100% (1)

where OD was the value of the cell on testing sample and OD0 was

the value of cells on tissue culture polystyrene (TCPS).

Figure 1. The 1H NMR spectra of PMA (A), mPEG2000-g-PMA (B)and hydrogel (C). The solvents were D2O for A and B, and DMSO-d6 for C.

Results and Discussion

In previous research,[24,26] poly(a,b-malic acid) was synthe-

sized by polycomdensation. mPEG-g-PMA copolymers with

two series of mPEG (Mw ¼ 750 and 2 000) and three different

graft degrees (20, 40 and 100%) were prepared. a-CDs were

threaded onto the grafted mPEG chains; the polyrotaxanes

were formed through host-guest interactions. Herein, we

used this method to fabricate supramolecular injectable

hydrogels. We found that only the copolymers with a 20%

graft degree could form hydrogels in both mPEG750 and

mPEG2000 grafted copolymers series. The compositions of

hydrogels are shown in Table 1.

The 1H NMR spectra of PMA, mPEG-g-PMA copolymer

and hydrogel are presented in Figure 1. The multi-peaks at

d¼ 3.0–3.2 were assigned to the protons of CH2 (b) in both a

and b-type units in poly(a,b-malic acid), the doublets at

d¼ 5.5 and 5.6 were attributed to the protons of CH (a) in

poly(a,b-malic acid). The chemical environment of the

random aggregated a and b-type repeated units in PMA

backbones was little different and thus led to the signals

splitting into multipeaks. Figure 1(B) shows the 1H NMR

spectrum of mPEG-g-PMA copolymer. In comparison with

the spectrum in Figure 1(A), the peaks at d¼ 3.4 and 3.6–3.8

ol. Biosci. 2009, 9, 1169–1175

WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

were attributed to the protons of OCH3 (d) and OCH2CH2 (c)

in mPEG. From the intensity ratio between CH3 (d) in mPEG

and CH (a) in poly(a,b-malic acid), the graft degree was

calculated. The 1H MNR spectrum of the hydrogel is

presented in Figure 1(C). According to the assignments of

the protons for polyrotaxanes reported in a previous

study,[13] the signal at d¼ 3.5 was the protons of OCH2CH2

in mPEG; it shifted from d¼ 3.75 (in Figure 1(B)) to d¼ 3.5.

The mole ratio of ethylene glycol (EG)/a-CD was calculated

from the integrities of OCH2CH2 (c) in mPEG and CH (1) in

a-CDs.

The GPC spectra of both PMA and mPEG-g-PMA were

multiply peaked, as shown in Figure 2. The multiple minor

peaks in the GPC curves were caused by polymers with low

molecular weight and impurities in the solvent. We selected

the main peak to calculate the relative molecular weight

and polydispersity. The Mn, Mw and polydispersity of PMA

www.mbs-journal.de 1171

B. He et al.

121086420

mPEG-g-PMA

PMA

Time (min)

Figure 2. The GPC spectra for PMA and mPEG2000-g-PMA.

1007550250-25-50

Hydrogel

mPEG-g-PMA

Endo

ther

mic

Temperature ( oC)

Figure 4. The DSC spectra of mPEG2000-g-PMAcopolymer andthe corresponding hydrogel (entry 4).

5040302010

Hydrogel

mPEG-g-PMA

2 Theta (degree)

Figure 5. The XRD spectra of mPEG2000-g-PMA copolymer andthe corresponding hydrogel (entry 4).

1172

were 2 870, 3 570 and 1.24, respectively. Those of mPEG-g-

PMA copolymer were 6 730, 8 940 and 1.33. The yields of the

polycondensation (after precipitation in diethyl ether) and

graft reaction were 57% and 90%, respectively.

The in situ gelation of the hydrogels is shown in Figure 3.

a-CDs, mPEG-g-PMA copolymers and DOX were dissolved in

water and mixed together. In order to discover the influence

of a-CD concentration on the gelation rate, hydrogels with

compositions according to entry 1 and 2 (Table 1) were used.

At the beginning, the mixtures were solutions, which could

flow in the bottles. The gelation of sample B was observed

3 min later. Sample A began to gelate after 5 min. The

gelation was complete after 12 min. This shows that a high

concentration of a-CDs accelerates the gelation process.

As mPEG2000 is a semi-crystalline polymer and

mPEG750 is a viscous liquid, the copolymer grafted

mPEG2000 was selected for comparison by DSC, XRD and

TGA. DSC and XRD were used to discover the structure of the

mPEG-g-PMA copolymer and hydrogel. The hydrogel was

freeze-dried before testing. There was only one endother-

mic peak in the spectrum of the mPEG-g-PMA copolymer at

about 45 8C, which was attributed to the melting of mPEG

crystals (Figure 4). No endothermic peak was observed in

the spectrum of the hydrogel; it revealed that the crystal of

mPEG was destroyed by a-CDs.

Figure 3. The in situ gelation process of hydrogels with different comp1; B: entry 2.

Macromol. Biosci. 2009, 9, 1169–1175


In the XRD spectra (Figure 5), two strong crystalline peaks

of mPEG appeared at 2u¼ 19.38 and 23.48 in mPEG-g-PMA

copolymer. While in the hydrogel, the two peaks vanished

and a new strong peak at 2u¼ 19.78 was presented. This is

the characteristic peak associated with a channel-type

crystalline structure in polyrotaxanes.[27–30] The XRD

spectrum implied that the a-CDs were threaded on mPEG

ositions, A: entry

chains and stacked polyrotaxane crystal.

The thermal decomposition of mPEG-g-

PMA copolymer and the corresponding

freeze-dried hydrogel was also studied

(Figure 6). There were two decomposition

steps in the mPEG-g-PMA spectrum

(Table 2). The first step, in which the

temperature of maximum decomposi-

tion rate and weight loss was 181.8 8Cand 15.64%, corresponded to PMA

DOI: 10.1002/mabi.200900225


5004003002001000

20

40

60

80

100

Hydrogel

mPEG-g-PMA

Wei

ght r

emai

ning

(%)

Temperature ( oC)

Figure 6. TGA spectra of mPEG2000-g-PMA copolymer and thecorresponding freeze-dried hydrogel (entry 4).

Scheme 1. The schematic formation mechanism of the supramo-lecular hydrogel.

Figure 7. SEM images of hydrgel with different amount of a-CDs,A1: hydrogel 1 (entry 1), � 100; A2: enlarged image of hydrogel 1,� 500; B1: hydrogel 2 (entry 2), � 100; B2: enlarged image ofhydrogel 2, � 500. The white arrows show the clusters.

decomposition. The second one was the decomposition of

mPEG. The temperature and weight loss were 391.5 8C and

80.99%. Three steps were observed in the TGA spectrum of

the freeze-dried hydrogel. The first step lasted from room

temperature to about 120 8C, which was attributed to the

evaporation of water absorbed by hydrophilic mPEG and a-

CDs. The decomposition temperatures of a-CDs and mPEG

were 293.4 and 387.5 8C, respectively. In the hydrogel

sample, the decomposition of PMA was not obvious; this

was because the content of PMA in the hydrogel was very

low. The decomposition of PMA was partially overlapped by

a-CD in step 2. The weight losses of step 2 and 3 were 49.76

and 24.04%, which were consistent with the compositions

of a-CDs and mPEG in the hydrogel.

According to the structure characterization results, the

schematic formation mechanism of the supramolecular

injectable hydrogel is proposed in Scheme 1. When a-CDs

and mPEG-g-PMA copolymers were mixed together in

aqueous solution, the a-CDs threaded onto mPEG

chains spontaneously to form polyrotaxanes. The poly-

rotaxanes self-assembled and stacked in crystal clusters.

The clusters crosslinked the water soluble polymeric

chains and the hydrogel was formed. If the concentration

of polyrotaxanes was high, the crystal clusters would

Table 2. The thermal decomposition of mPEG-g-PMA copolymer (mP

Sample Step 1

Weight Loss Temp.a) Weig

% -C

mPEG-g-PMA 15.64 181.8 8

Hydrogel 4.04 – 4

a)The temperature of maximum decomposition rate.

Macromol. Biosci. 2009, 9, 1169–1175


precipitate because of the strong hydrogen bonding

interaction inside; this is why the copolymers with higher

mPEG graft degrees could not be fabricated into hydrogels.

The SEM was used to observe the gelated hydrogels in

Figure 3. The freeze-dried hydrogels were porous networks

and the surface morphologies were closely related to the

amount of a-CDs (Figure 7). The pores of hydrogel 1 were

homogeneous and the average pore size was around 30mm.

The surface morphology of hydrogel 2 was very different;

EG, Mw ¼ 2 000) and freeze-dried hydrogel (entry 4).

Step 2 Step 3

ht Loss Temp. a) Weight Loss Temp.a)

% -C % -C

0.99 391.5 – –

9.76 293.4 24.04 387.5


B. He et al.

Figure 8. The release of DOX from hydrogel 1, 2, 3 and 4.

1174

there were some clusters on the surface, and the average

pore size was smaller than that of hydrogel 1. According to

Figure 9. Cancer cell (U87MG) morphologies coculctured with blank hydrogel andhydrogel loaded with doxorubicin: A1: TCPS for 24 h; A2: TCPS for 48 h; A3: TCPS for72 h; B1: Blank hydrogel for 24 h; B2: Blank hydrogel for 48 h; B3: Blank hydrogel for 72 h;C1: Hydrogel with drug for 24 h; C2: Hydrogel with drug for 48 h; C3: Hydrogel with drugfor 72 h. The amount and the composition of blank hydrogel and hydrogel loaded withdrug were according to entry 4 in Table 1.

previously reported stoichiometry of

EG/CD in polyrotaxanes,[26] nearly all

the a-CDs in hydrogel 1 were threaded

onto mPEG chains. The amount of a-CDs

was excessive in hydrogel 2, and the

unthreaded a-CDs formed aggregated

clusters in the hydrogel.

The in vitro drug release was carried

out in PBS (Figure 8). The burst release

phenomenon, which commonly exists in

drug delivery systems, was not serious in

these supramolecular injectable hydro-

gels. The factor to affect the drug release

rate greatly was the composition, espe-

cially the content of a-CDs. Generally,

higher contents of mPEG-g-PMA copoly-

mer and a-CDs in hydrogels resulted in a

longer release time. The concentration of

a-CDs in hydrogel 2 was higher than that

in hydrogel 1, and its drug release rate

was slower than that of hydrogel 1. The

sustaining release time of hydrogel 1 was

about 90 h and that of hydrogel 2 was

more than 180 h. The compact pores and

smaller pore size of hydrogel 2 led to

slower drug diffusion and thus led to

longer sustaining release time. Compared

with the drug release of hydrogel 2 and 4,

it could be found that though the total

content of mPEG-g-PMA copolymers and

a-CDs in hydrogel 2 was lower than that

in hydrogel 4, its release rate was slower

than that of hydrogel 4. This is maybe due

Macromol. Biosci. 2009, 9, 1169–1175


to two factors, one was the chain length of mPEG, and the

other was the EG/a-CD ratio.[10,22]

The anti-tumor effect of the released DOX was evaluated

in vitro. Figure 9 shows the morphologies of glioma cancer

cells (U87MG) co-cultured with hydrogels. Series A was the

TCPS control, series B was the blank hydrogel, and series C

was the hydrogel loaded with DOX. The attachment,

spreading and morphology of the cancer cells in series A

and B were normal, and the cell proliferation was obvious

within the 3 co-culture days. In series C, it was very

different. On the first day, the quantity of attached cells

was lower than that in series A or B, and some detached

cells were observed. On the second day, the cell spreading

was suppressed, many cells shrank and more detached cells

were observed. It became worse on the third day, when

nearly all the cancer cells detached and shrank to global

shape. The variation of cells morphologies demonstrated

that the anti-tumor effect of the supramolecular injectable

hydrogel drug delivery system was positive.

The anti-tumor effect of the released DOX was evaluated

quantitatively by cell viability. The test result is presented

DOI: 10.1002/mabi.200900225


0

20

40

60

80

100

120

724832248

Cel

l via

bilit

y (%

)

Time (h)

Hydrogel 4 Hydrogel 4 +DOX

Figure 10. Cell viability of blank hydrogel and hydrogel loadedwith DOX. The amount and the composition of blank hydrogeland hydrogel loaded with drug were according to entry 4 inTable 1.

in Figure 10. The cancer cells in both the TCPS control and

blank hydrogel kept high viability. The cell viability in blank

hydrogel decreased by about 15% within the first 2 d and it

recovered subsequently. This is due to the weak acidity of

the blank hydrogel. The cell viability recovery was because

of the neutralization of the cell culture medium. The cell

viability of the hydrogel loaded with DOX decreased

continuously; it was less than 40% after 3 d. The sustaining

released drugs killed cancer cells or suppressed the growth

of cancer cells and thus led to the cell viability decreasing.

Conclusion

A supramolecular injectable hydrogel composed of a-CDs

and mPEG-g-PMA copolymers was fabricated. The char-

acterization results showed that the driving force for in situ

gelation was the host-guest interaction between a-CDs and

mPEG chains. The polyrotaxane crystals acted as physical

crosslink sites. The anti-tumor drug doxorubicin was loaded

in the hydrogel. The drug release study revealed that the

burst release was restrained; the sustaining drug release

time could last more than 3 d. The anti-tumor effect of the

released DOX was evaluated in vitro. The released DOX

could kill cancer cells and/or suppress the growth of cancer

cells effectively. This injectable hydrogel is promising for

minimally invasive therapeutic drug delivery.

Acknowledgements: This research work was supported by theNational Basic Research Program of China (National 973 program,No. 2005 CB623903), National Science Foundation of China (NSFC,No. 20604016, 50830105), Sichuan Youth Science and Technology

Macromol. Biosci. 2009, 9, 1169–1175


Foundation (No. 07ZQ026-013) and The Scientific ResearchFoundation for Returned Overseas Chinese Scholars, State Educa-tion Ministry (No. 20071108-18-5). The authors gratefullyacknowledge Prof. Hua Ai and Dr. Gang Liu for their kindnessin providing U87MG cancer cells.

Received: June 24, 2009; Revised: September 22, 2009; Publishedonline: November 12, 2009; DOI: 10.1002/mabi.200900225

Keywords: biodegradable; drug delivery systems; gelation;hydrogels; supramolecular structures

[1] A. Patel, K. Mequanint, Macromol. Biosci. 2007, 7, 727.[2] J. Kopecek, J. Y. Yang, Polym. Int. 2007, 56, 1078.[3] B. H. Lee, B. Vernon, Macromol. Biosci. 2005, 5, 629.[4] S. Kumar, K. J. Himmelstein, J. Pharm. Sci. 1995, 84, 344.[5] Q. P. Hou, D. Y. S. Chau, C. Pratoomsoot, P. J. Tighe, H. S. Dua,

K. M. Shakesheff, J. Pharm. Sci. 2008, 97, 3972.[6] C. L. He, S. W. Kim, D. S. Lee, J. Controlled Release 2008, 127,

189.[7] J. T. Zhang, S. W. Huang, J. Liu, R. X. Zhuo, Macromol. Biosci.

2005, 3, 192.[8] H. S. Choi, K. Yamamoto, T. Ooya, N. Yui, ChemPhysChem

2005, 6, 1081.[9] I. Tomatsu, A. Hashidzume, A. Harada, Macromol. Rapid

Commun. 2006, 27, 238.[10] J. Li, X. J. Loh, Adv. Drug Deliver. Rev. 2008, 60, 1000.[11] P. Lo Nostro, J. R. Lopes, B. W. Ninham, P. J. Baglioni, J. Phys.

Chem. B 2002, 106, 2166.[12] C. Cheng, R. Tang, F. Xi, Macromol. Rapid Commun. 2005, 26,

744.[13] J. Li, A. Harada, M. Kamachi, Polym. J. 1994, 26, 1019.[14] S. Zhao, L. Zhang, D. Ma, J. Phys. Chem. B 2006, 110, 12225.[15] J. Li, X. P. Ni, K. W. Leong, J. Biomed, Mater. Res., Part A 2003,

65A, 196.[16] J. Li, X. Li, Z. Zhou, X. Ni, K. W. Leong, Macromolecules 2001, 34,

7236.[17] J. Li, X. Li, X. Wang, X. Ni, H. Li, K. W. Leong, Biomaterials 2006,

27, 4132.[18] K. M. Huh, T. Ooya, W. K. Lee, S. Sasaki, I. C. Kwon, S. Y. Jeong,

N. Yui, Macromolecules 2001, 34, 8657.[19] K. M. Huh, Y. W. Cho, H. Chung, I. C. Kwon, S. Y. Jeong, T. Ooya,

W. K. Lee, S. Sasaki, N. Yui, Macromol. Biosci. 2004, 4, 92.[20] L. He, J. Huang, Y. Chen, X. Xu, L. Liu, Macromolecules 2005, 38,

3845.[21] E. Sabadini, T. Cosgrove, Langmuir 2003, 19, 9680.[22] X. Li, J. Li, J. Biomed. Mater. Res., Part A 2008, 86A, 1055.[23] L. X. Ren, L. H. He, T. S. Sun, X. Dong, Y. M. Chen, J. Huang, C.

Wang, Macromol. Biosci. 2009, 10.1002/mabi.200900225.[24] B. He, E. Wan, M. B. Chan-Park, Chem. Mater. 2006, 18, 3946.[25] J. M. Gao, J. Ming, B. He, Y. J. Fan, Z. W. Gu, X. D. Zhang, Eur. J.

Pharm. Sci. 2008, 34, 85.[26] J. Zeng, Y. Li, J. J. Li, R. Wang, B. He, Y. Nie, X. L. Luo, Z. R. Zhang,

Z. W. Gu, Chinese J. Polym. Sci. 2009, 27(5), 729.[27] I. N. Topchieva, A. E. Tonelli, I. G. Panova, E. V. Matuchina, F. A.

Kalashnikov, V. I. Gerasimov, C. C. Rusa, M. Rusa, M. A. Hunt,Langmuir 2004, 20, 9036.

[28] A. Harada, M. Kamachi, Macromolecules 1990, 23, 2821.[29] J. Li, X. Ni, Z. Zhou, K. W. Leong, J. Am. Chem. Soc. 2003, 125,

1788.[30] T. Uyar, M. A. Hunt, H. S. Gracz, A. E. Tonelli, Cryst. Growth Des.

2006, 6, 1113.


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