Light-responsive amphiphilic copolymer coated nanoparticles as nanocarriers and real-time monitors for controlled drug release

Article abstract of DOI:10.1039/c3tb21269f

Herein, light-responsive nanocarriers based on hollow mesoporous silica (HMS) nanoparticles modified with spiropyran-containing light-responsive copolymer (PRMS-FA) were fabricated via a simple self-assembly process. HMS modified with long-chain hydrocarb

Full text of DOI:10.1039/c3tb21269f

Journal of
Materials Chemistry B
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PAPER
Light-responsive amphiphilic copolymer coated
nanoparticles as nanocarriers and real-time
monitors for controlled drug release†
Cite this: J. Mater. Chem. B, 2014, 2,
1182
a
ab
*
Qingjian Xing, Najun Li,
Dongyun Chen,a Wenwei Sha,a Yang Jiao,c Xiuxiu Qi,a
Qingfeng Xuab and Jianmei Lu
ab
*
Herein, light-responsive nanocarriers based on hollow mesoporous silica (HMS) nanoparticles modified
with spiropyran-containing light-responsive copolymer (PRMS-FA) were fabricated via a simple self-
assembly process. HMS modified with long-chain hydrocarbon octadecyltrimethoxysilane was an ideal
base material owing to its good biocompatibility and drug capability. The spiropyran-containing
amphiphilic copolymer could shift its hydrophilic–hydrophobic balance to become hydrophilic upon UV
(l ¼ 365 nm) irradiation and then break away from the hydrophobic surface of the HMS core, followed
by the uncaging and release of the pre-loaded anticancer drug. Simultaneously, the fluorescence
resonance energy transfer (FRET) process based on the structural transformation of PRMS-FA was
observed, which could act as a real-time monitor for the light-controlled drug release. Our model
experiments in vitro tested and verified that this composite nanocarrier has good biocompatibility, active
tumour targeting to the folate receptor over-expressed in tumour cells, is non-toxic to normal cells and
that light-controlled drug release with real-time monitoring can be achieved.
Received 16th September 2013
Accepted 22nd November 2013
DOI: 10.1039/c3tb21269f
www.rsc.org/MaterialsB
o-nitrobenzyl and coumarin.6 In addition, such light-responsive
drug carriers could be modied by introduction of tumour
Introduction
targeting compounds to the polymers, such as folic acid,
peptide and antibodies, to achieve active targeting ability.7
Hence light irradiation would disaggregate the light-responsive
drug carriers to release the loaded drugs aer they had gathered
at the tumour site.8
Spiropyran, one of the photochromic compounds, exists as a
closed-loop non-colored spiroform (SP) under visible light
irradiation, whereas it can be converted to an open-loop colored
merocyanine form (MC) under UV irradiation (l ¼ 365 nm).
What is more, the SP form is hydrophobic and can be converted
into the hydrophilic MC form by UV irradiation.9 So spiropyran-
containing amphiphilic polymeric micelles may disaggregate
owing to the structural transformation upon UV irradiation,
resulting in the release of drugs pre-loaded in the micelles. This
could provide a strategy to design a spiropyran-based light-
responsive nanocarrier for light-controlled drug delivery and
release.
In recent years, nanoscale drug delivery systems (NDDS) have
attracted considerable attention for their potential applications
in biomedical elds.1 Among them, polymeric micelles assem-
bled by amphiphilic copolymers containing stimuli-responsive
moieties which respond to various external stimuli, such as
light, pH and temperature, have been introduced as drug
carriers for controlled release.2 Light, having higher selectivity
in time and space than other stimuli, is considered to be an
ideal trigger in NDDS.3 Some light triggered NDDS based on
functional nanoparticles, upconverting nanoparticles for
example, have been reported.4 Since good control over the
reaction could be ensured by adjusting the wavelength and
intensity of the light,5 light-triggered drug delivery systems
based on polymeric micelles have been designed to be applied
in controlled drug release through introduction of photo-
chromic compounds such as azobenzene, spiropyran,
In controlled drug release systems, real-time monitoring of
the drug release plays an important role. The uorescence
resonance energy transfer (FRET) process was conrmed to act
as a monitor in drug nanocarriers based on the energy transfer
between coumarin and FITC.10 Such a strategy was also suitable
for application in a light-controlled drug release system. Spi-
ropyran, could act as the acceptor in the FRET process while
rhodamine B, a common uorescent compound, can be used as
the donor owing to their well matching absorption and
aJiangsu Key Laboratory of Advanced Functional Polymer Design and Application,
College of Chemistry, Chemical Engineering and Materials Science, Soochow
University, Suzhou, 215123, China. E-mail: linajun@suda.edu.cn; lujm@suda.edu.
cn; Fax: +86-512-6588-0367; Tel: +86-512-6588-0367
bState Key Laboratory of Treatments and Recycling for Organic Effluents by Adsorption
in Petroleum and Chemical Industry, Suzhou, 215123, China
cSchool of Radiation Medicine and Protection, Medical College of Soochow University,
Suzhou, Jiangsu 215123, China
† Electronic supplementary information (ESI) available: Experimental section. See
DOI: 10.1039/c3tb21269f
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emission bands.11 In the FRET process, the open-loop MC form targeting copolymer PRMS-FA via a simple self-assembly
of spiropyran has a strong overlap and quenches the uores- process. Hence the core–shell light-triggered nanocomposites
cence of rhodamine B upon UV irradiation while the SP form for targeting drug delivery and controlled release were
does not. Therefore, it is possible that an amphiphilic copol- obtained. We mainly investigated light controlled drug
ymer containing both spiropyran and rhodamine B moieties release under UV (l ¼ 365 nm) irradiation and the possibility
can show a similar FRET process aer UV irradiation. Since the of the uorescence resonance energy transfer (FRET) process
FRET process occurred simultaneously with the disaggregation being used as the real-time monitor for pre-loaded drug
of polymeric micelles upon UV irradiation, it could act as a real- release.
time monitor for light-triggered drug release based on changes
in the uorescence signal.
To improve the drug-loading capability and morphological
stability of polymeric micelles, some functional inorganic
nanoparticles were incorporated into the micelles by a self-
Experimental section
Materials
assembly process.12 This strategy could combine the benecial
properties of the inorganic nanoparticles with polymeric
micelles for better application in drug delivery.13 Among such
inorganic nanoparticles, hollow mesoporous silica (HMS)
nanoparticles are ideal base materials with high biocompati-
bility and drug loading capability.14 Besides, the nanoparticles
can be easily modied owing to the large amount of Si–OH on
the outer surface. To achieve a “smart” controlled drug release
system, HMS nanoparticles are usually coated with stimuli-
responsive polymer and such composite construction was
conrmed to be stable without external stimulus according to
our previous work.15,16
2,3,3-Trimethyl-3H-indol, 2-hydroxy-5-nitrobenzaldehyde and 2-
bromoethanol were purchased from TCI. (2-(Acryloyloxy) ethyl)
trimethylammonium chloride (AETAC, 80 wt% in water), 2,20-
azobis(2-methylpropionamidine)
>97.0%), octadecyltrimethoxysilane
dihydrochloride
(C18),
(V-50,
hexadecyl-
trimethylammonium bromide (CTAB, >99%) and poly(ethylene
glycol) methacrylate (MAPEG, Mn average ¼ 526) were all
purchased from Aldrich. Folic acid, methacryloyl chloride,
methanesulfonyl chloride, sodium azide, triphenylphosphine
were all purchased from Shanghai Chemical Reagent Co. Ltd.
Rhodamine B, N-hydroxysuccinimide (NHS, 99%) and
dicyclohexacarbodiimide (DCC, 99%) were purchased from Alfa
Aesar.
Here we designed a spiropyran-containing light-responsive
amphiphilic copolymer which was coated onto the surface of
HMS nanoparticles forming a drug nanocarrier (Scheme 1).
Firstly, a spiropyran-containing amphiphilic copolymer, poly-
(RBM-MAPEG-SPMA) (PRMS), was synthesized by radical
copolymerization. Poly(ethylene glycol) (PEG), a popular
hydrophilic polymer with good biocompability and stealth
behaviour, was introduced as the hydrophilic part. Folic acid
(FA), a common tumour targeting compound for tumour cells
with over-expressed folate receptors (FR(+)), was then conju-
gated with the PRMS to achieve the nal multifunctional
copolymer, PRMS-FA. Hollow mesoporous silica (HMS)
nanoparticles with high drug loading capacity acted as the
drug delivery matrix to be modied with the light-responsive
Modication of HMS with C18 (HMS@C18)
First, the hollow mesoporous silica (HMS) nanoparticles were
synthesized according to the literature with some modica-
tions.15 Then, 100 mg HMS nanoparticles were dispersed in
20 mL anhydrous acetonitrile before 5 mL C18 was added. The
obtained suspension was stirred for 24 h and collected by
centrifugation, washed with acetonitrile and ethanol several
times, and dried under vacuum for further usage.
Synthesis of hydrophobic monomer 10-(2-methacryloxyethyl)-
30,30-dimethyl-6-nitro-spiro(2H-1-benzopyran-20,20-indoline)
(SPMA)
The synthetic route to SPMA is shown in Scheme S1.† SPOH was
synthesized according to the literature.17 The general procedure
employed for the preparation of SPMA is as follows. SPOH (1.0
g, 2.84 mmol) was dissolved in dry CH2Cl2 (15 mL) and Et3N
(868 mL, 6.25 mmol) was added. The mixture was cooled to 0 ꢀC
under argon and protected against exposure to visible light.
Methacryloyl chloride (550 mL, 5.68 mmol) in dry CH2Cl2 (5 mL)
was then added dropwise over 0.5 h. The mixture was allowed to
stir over night at room temperature. Aer the solvent was
removed using a rotary evaporator, the residues were puried by
silica gel column chromatography (petroleum ether–CH2Cl2 ¼
1 : 2) to give SPMA as a green solid (0.48 g, 40%).
1H NMR (400 MHz, CDCl3), d (ppm): 7.97–8.05 (m, 2H), 7.17–
7.24 (m, 1H), 7.09 (d, 1H), 6.89 (q, 2H), 6.74 (q, 2H), 6.07 (d, 2H),
5.87 (d, 1H), 5.56 (d, 2H), 4.3 (t, 2H), 3.37–3.62 (m, 2H), 1.91 (s,
3H), 1.28 (s, 3H), 1.16 (s, 3H).
Scheme 1 Schematic depiction of light controlled drug release from
HMS/C18/PRMS-FA and the FRET process.
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Synthesis of rhodamine B (2-hydroxyethyl acrylate) ester
(RBM)
Preparation of HMS@C18@PRMS-FA
HMS@C18 (5 mg) was dispersed in 1 mL tetrahydrofuran by
The general procedure is as follows. Rhodamine B (10 g) was sonication for 5 min. PRMS-FA (30 mg) was dissolved in the
dissolved in 1,2-dichloroethane (150 mL) and thionyl chloride solution. Then, 5 mL distilled water was added dropwise to the
ꢀ
(10 mL) was added slowly to the solution below 0 C. Then the above solution with vigorous stirring and the resulting mixture
reaction mixture was stirred under reux for 8 h. Aer cooling was stirred vigorously over night while the tetrahydrofuran
down to room temperature, 2-hydroxyethyl acrylate (15 mL) was evaporated. Aer that, the product was separated by centrifu-
added to the solution and the mixture was allowed to stir gation and washed with distilled water several times.
overnight. The resulting product was puried by silica gel
column chromatography (CH2Cl2–CH3OH ¼ 30 : 1, v/v) to give
Drug loading and release experiment
RBM as a purple solid (510 mg, 42.2%).
1H NMR (400 MHz, CDCl3), d (ppm): 8.32 (d, 1H), 7.85 (t, 1H),
7.76 (d, 1H), 7.33 (d, 1H), 7.07 (d, 2H), 6.94 (d, 2H), 6.81 (s, 2H),
To evaluate the drug loading and release properties, doxoru-
bicin (DOX) was used as a model anticancer agent. DOX was
extracted from doxorubicin hydrochloride (DOX$HCl) accord-
6,37 (m, 1H), 6.05 (m, 1H), 5.87 (m, 1H), 4.29 (m, 2H), 4.17 (m,
ing to the procedure reported previously.19 HMS/C18 (8, 4 and
2H), 3.66 (m, 8H), 1.34 (t, 12H).
0.8 mg mLꢁ1) was dispersed into 1 mL phosphate buffered
saline (PBS) solution, then DOX solution (5 mg mLꢁ1, 80 mL) was
Synthesis of MAPEG amine (MAPEG-NH2)
added. The mixture was stirred for 24 h (protected against
visible light exposure). Aerwards, DOX/HMS/C18 were
obtained by centrifugation and washed several times, then
MAPEG amine was synthesized according to the literature.18
MAPEG azide was prepared in two steps (see ESI†).
Then MAPEG azide (2.76 g, 5 mmol) was dissolved in THF
(5 mL). Triphenylphosphine (1.572 g, 6 mmol) was added, N2
evolution was observed, and the mixture was stirred at room
temperature for 2 h. Distilled water (0.5 mL) was added, and the
mixture was stirred at room temperature overnight. The mixture
was evaporated and dissolved in CH2Cl2. The solution was
washed with 1 N HCl (twice). The aqueous layer was made basic
with 1 N NaOH and extracted with CH2Cl2 (twice). The organic
layer was dried over MgSO4, ltered, and evaporated to give
MAPEG amine as a yellow liquid (1.32 g, 50%).
ꢀ
dried at 60 C under vacuum.
To evaluate the amount of drug loaded into HMS/C18, UV–
Vis spectroscopy was used for analysis. First, the calibration
curve of DOX was determined by taking absorbance vs. DOX
concentration between 0 and 5 mg Lꢁ1 as parameters, and the
calibration curve was tted to the Lambert–Beer law as follows:
A ¼ 0.0014 + 0.0091C
where A is the absorbance and C is the concentration (mg Lꢁ1).
Aer adsorption, the DOX solution (100 mL) was extracted and
diluted to 10 mL, and then analyzed by UV–Vis spectroscopy at a
wavelength of 485 nm. Drug loading content and drug loading
efficiency can be calculated as follows:
1H NMR (400 MHz, CDCl3), d (ppm): 6.12 (m, 1H), 5.57 (m,
1H), 4.29 (m, 2H), 3.74 (t, 2H), 3.64 (m, 32H), 3.54 (t, 2H), 1.94
(s, 3H).
Drug loading content (wt%) ¼ (weight of loaded drug/weight of
nanocomposites) ꢂ 100%
Synthesis of the light-responsive copolymer poly(RBM-
MAPEG-SPMA) (PRMS)
PRMS was synthesized by radical copolymerization. MAPEG
amine (315.6 mg, 0.6 mmol), SPMA (84 mg, 0.2 mmol) and RBM
(1.2 mg, 0.002 mmol) were dissolved in cyclohexanone (2 mL).
Then, AIBN (1.1 mg) was added to the mixture. The reaction
mixture was stirred at 70 ꢀC for 12 h under argon. Aer that, the
solution was suspended in anhydrous ether (100 mL) and the
copolymer was obtained aer centrifugation at 8000 rpm for
10 min. The copolymer was washed several times before being
dried under vacuum for further usage.
Drug loading efficiency (%) ¼ (weight of loading drug/weight of
drug in feed) ꢂ 100%
The drug release test was performed by suspending the DOX-
loaded nanoparticles in phosphate buffered saline (PBS) at
ꢀ
37 C. The solution was shaken vigorously under visible light
and UV (l ¼ 365 nm) irradiation alternately. To determine the
release amount at any given time, a portion of solution was
withdrawn aer centrifugation and the same volume PBS was
added to keep the volume constant. The drug concentration in
the withdrawn solution was analyzed by measuring the absor-
bance of DOX using UV–Vis spectroscopy.
Synthesis of folate conjugated copolymer PRMS-FA
The conjugation procedure was briey as follows. FA–NHS was
prepared at rst (see ESI†). Then FA–NHS (30 mg) was added to
dry pyridine containing 200 mg of the above copolymer. The
solution was stirred for 12 h at room temperature (protected
Cell culture and preparation
against visible light exposure). Then, the mixture was precipi- Human KB cell lines with folic acid receptors (FR(+)) and A549
tated in 10% acetone in anhydrous ether. The copolymer PMRS- human alveolar adenocarcinoma cell lines without folic acid
FA was obtained by centrifugation and stored in the dark receptors (FR(ꢁ)) were purchased from Shanghai Cell Institute
ꢀ
at 4 C.
Country Cell Bank, China and cultured as monolayers in
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RPMI-1640 medium supplemented with 10% heat-inactivated mLꢁ1. The medium was removed aer 3 h of incubation and the
fetal bovine serum at 37 ꢀC in a humidied incubator (5% CO2 cells were washed before treatment with various UV light doses.
in air, v/v).
Aer a further 1 h of incubation, the cells were washed and
observed using CLSM at a wavelength of 510 nm.
In vitro cytotoxicity
Characterization
The aminoxanthene dye, sulforhodamine B (SRB) was used as
an assay for assessing the effects of drug carriers
(HMS@C18@PRMS-FA) in various concentrations. Briey,
strongly growing KB cells were placed in 96-well plates (1.3 ꢂ
104 cells per well) and four duplicate wells were set up in the
sample. The culture medium was replaced with a medium
containing different concentrations of drug carriers and
cultured at 37 ꢀC in a humidied incubator (5% CO2 in air, v/v)
with the cells anchored to the plates. Aer being cultured for
72 h, the medium was poured away and 10% (w/v) trichloro-
acetic acid in Hank's balanced salt solution (100 mL) was added
1H NMR spectra were measured using an INOVA 400 MHz NMR
instrument. FT-IR measurements were performed using KBr
pellets on a Nicolet 4700 spectrometer (Thermo Fisher Scien-
tic) in the range of 400–4000 cmꢁ1. The UV–Vis absorption
spectra were measured on a TU-1901 spectrophotometer.
Fluorescence spectra were measured on a HORIBA Jobin Yvon's
uorescence spectrouorometer. Gel permeation chromatog-
raphy (GPC) analysis was carried out using a Waters 1515 pump
and a differential refractometer, THF was used as the mobile
phase at a ow rate of 1.0 mL minꢁ1. TEM images were obtained
using a TecnaiG220 electron microscope at an acceleration
voltage of 200 kV. Brunauer–Emmett–Teller (BET) and Barrett–
Joyner–Halenda (BJH) analyses were used to determine the
surface area, pore size and pore volume and measurements
were obtained with a Quantachrome Autosorb 1C apparatus at
ꢁ196 ꢀC under continuous adsorption conditions. The UV
(365 nm, 6 W) and light-emitting diode (LED) lamps (620 nm, 15
W) were used as light sources for UV and visible light irradia-
tion. Confocal laser scanning microscopy (CLSM) images were
observed by a confocal laser scanning microscopy (Olympus,
FV 1000).
ꢀ
and stored at 4 C for 1 h. Then the stationary liquid was dis-
carded, the cells were washed with deionized water ve times
before air drying and stained with 0.4% (w/v) SRB solution
(100 mL per well) for 30 min at room temperature. Aer
removing the SRB, the cells were washed with 0.1% acetic acid
solution ve times. Bound SRB dye was solubilized with
10 mmolꢁ1 Tris-base solution (150 mL, pH ¼ 10.5). The optical
density (OD) value of each individual well was calculated by
using a spectrophotometer to measure the absorbance at
531 nm.
The cell cytotoxicity of light-triggered DOX release was
assessed by the standard MTT assay. To test the cytotoxicity of
the DOX-loaded drug carriers with and without UV irradiation,
KB cells were seeded in a 96-well plate at a density of 1.3 ꢂ 104
cells per well and cultured in 5% CO2 at 37 ꢀC for 24 h. Then the
Results and discussion
Synthesis and FRET behaviour of PRMS-FA
drug carriers were added to the medium. For the drug carrier The amphiphilic copolymer PRMS-FA was synthesized
without UV irradiation, the cells were incubated for varying time following the procedure shown in Scheme 2. The GPC data for
the copolymers with different molar ratios is summarized in
Table S1.† PRMS (Mn ¼ 11 600, PDI ¼ 1.243) was chosen to be
conjugated with FA–NHS to obtain PRMS-FA (Mn ¼ 16 700, PDI
¼ 1.72). Approximately 63.7% of PEG chains on average were
durations. For the drug carrier with UV irradiation, the carrier
was removed aer 3 h of incubation followed by exposure to UV
light. Then the cells were incubated for another 24 h in the dark.
Aer that, the medium was removed before 100 mL MTT solu-
tion (0.5 mg mLꢁ1) was added and incubated for a further 4 h.
The medium was then replaced with 100 mL DMSO per well and
the absorbance was monitored using a microplate reader at a
wavelength of 485 nm. The cytotoxicity was expressed as the
percentage of cell viability compared to the untreated cells.
CLSM observation of the cell uptake and the FRET process
The nanocarrier is easily observed, owing to the uorescent
rhodamine B component, by confocal laser uorescence
microscopy (CLSM). For CLSM observation of the cell uptake,
A549 cells and KB cells were seeded in 96-well plates (1.3 ꢂ 104
cells per well) and incubated overnight at 37 ꢀC in a humidied
incubator. The sample of DOX/HMS/C18/PRMS-FA was
dispersed in RPMI-1640 medium. Cells were cultured with the
DOX/HMS/C18/PRMS-FA solution for
a certain time and
observed using CLSM aer washing three times with PBS. For
CLSM observation of the FRET process accompanied by DOX
release under UV irradiation, the KB cells were treated with
DOX/HMS/C18/PRMS-FA at the same concentration of 500 mg Scheme 2 Schematic synthetic route to PRMS-FA.
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conjugated with folic acid according to the incremental quan- non-colored spiropyran in SP form turned into the colored MC
tity of the molecular weight. The structure of PRMS-FA was form, which resulted in the color change of copolymer solution
conrmed by 1H NMR (see Fig. S1†), in which the characteristic to a much darker colour. Meanwhile, the uorescence emission
peaks in SPMA (d ¼ 4.3 and 6.75–7.0 ppm), RBM (d ¼ 8.10– of the copolymer at 510 nm decreased signicantly, which could
8.16 ppm) and folic acid (d ¼ 6.64, 7.65 and 8.64 ppm) were all be ascribed to the quenching of RBM uorescence by the MC
observed correspondingly, conrming that the amphiphilic form of the SPMA moieties when the FRET process occurred.
copolymers PRMS-FA were synthesized successfully.
Since the amphiphilic copolymer PRMS-FA would become
Spiropyran can be reversibly switched between the non- hydrophilic upon UV irradiation, the accompanied FRET
colored closed-loop state (SP, spiroform) and another colored process in the form of uorescence quenching could provide
open-loop state (MC, merocyanine form) by irradiation with evidence for such a transformation.
visible light and ultraviolet light. The UV–Vis absorption of
The surface of HMS was enriched with Si–OH, which could
spiropyran methacrylate (SPMA) before and aer irradiation be easily modied with C18. The FT-IR spectra of HMS, HMS/
with UV light at 365 nm is shown in Fig. 1a. It is worth noting C18 and DOX/HMS/C18 are shown in Fig. 2. The Si–OH bond
that the absorption bands of SPMA match well with the emis- (950 cmꢁ1) in the spectrum of HMS suggested that a large
sion bands of rhodamine B (2-hydroxyethyl acrylate) ester number of silanol groups existed on the surface of HMS for
(RBM) under UV irradiation. The FRET process requires overlap coupling long alkyl chains and functional molecules. As shown
of the emission bands of the donor and acceptor in general. in the spectrum of HMS/C18, a strong band of C–Hx (2890 cmꢁ1
)
Therefore, RBM could act as the donor while SPMA acts as the revealed the successful modication of alkyl chains on the
acceptor in this system. Thus, a similar FRET process is HMS. Compared to the spectrum of HMS/C18, a new C]O bond
expected to occur along with the transformation of PRMS-FA (1740 cmꢁ1) appeared in the spectrum of DOX/HMS/C18, sug-
from an amphiphilic copolymer to a hydrophilic polymer upon gesting that DOX molecules were loaded into the nanoparticles
UV irradiation. The photographs and the change in uores- successfully.
cence emission intensity under alternating UV and visible light
irradiation shown in Fig. 1b demonstrate the process. Under
Assembly of HMS/C18/PRMS-FA
visible irradiation, the bright red solution of PRMS-FA showed
quite high uorescence intensity (excited at 510 nm) which was
derived from the rhodamine B moiety. Upon UV irradiation, the
The folate conjugated amphiphilic copolymer PRMS-FA was
coated onto HMS nanoparticles via self-assembly. Aer modi-
cation with C18 alkyl chains, the surface of HMS nanoparticles
became hydrophobic. The hydrophobic part of the PRMS-FA
associates with the hydrophobic C18 alkyl chains through
hydrophobic van der Waals interactions during the self-
assembly process, as a result the amphiphilic copolymer PRMS-
FA was coated onto the surface of the HMS nanoparticles
successfully. The detailed morphological features of the nano-
particles were determined by TEM.
As shown in Fig. 3, aer the pure HMS was modied with
C18, no signicant changes were observed and the pores in the
silica shell remained clearly visible, suggesting that the modi-
cation with C18 does not affect the porosity of HMS. Aer self-
assembly with PRMS-FA, the prole of HMS/C18/PRMS-FA
became blurred and a thick layer of polymer lm could be
clearly observed. Aer UV (l ¼ 365 nm) irradiation, the
Fig. 1 (a) Superposition of the fluorescence emission of rhodamine B
(2-hydroxyethyl acrylate) ester (RBM) and the absorbance of SPMA
before (SP form) and after (MC form) UV irradiation (the inset images
show photographs of the RBM solution (1), mixed solution of RBM and
SPMA before (2) and after (3) UV irradiation); (b) fluorescence spectra
of the light-responsive copolymer PRMS-FA before and after UV
(l ¼ 365 nm) irradiation (the inset images show photographs of PRMS-
FA solution under visible light irradiation (1) and UV irradiation (2)).
Fig. 2 FTIR spectra of HMS, HMS/C18, DOX and DOX/HMS/C18.
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Fig. 4 Release of DOX in vitro from HMS/C18/PRMS-FA with and
without light irradiation at 37 ꢀC.
Fig. 3 TEM images of HMS (a), HMS/C18 (b), HMS/C18/PRMS-FA
(c) and HMS/C18/PRMS-FA upon UV irradiation (d).
The results of drug release from HMS/C18/PRMS-FA in vitro
are shown in Fig. 4. DOX-loaded HMS/C18/PRMS-FA was rst
immersed in PBS at 37 ꢀC. Upon visible light irradiation for 8 h,
less than 10 wt% of DOX was released. There was almost no
further release aer another 4 h irradiation. However, a fast and
high release of DOX was measured aer the nanocarrier was
irradiated with UV light (l ¼ 365 nm). No further release of DOX
was measured aer visible light irradiation for 4 h. The process
of drug release continued upon UV irradiation and approxi-
mately 70 wt% of DOX was released within 100 h. There was
negligible release of DOX when the nanocarrier was placed in
the dark. These results conrmed that DOX/HMS/C18/PRMS-FA
would only release drugs aer irradiation with UV light.
copolymer layer disappeared which conrmed our previous
thought. Furthermore, the DLS result (Fig. S3†) showed that the
nanoparticles gained an average diameter of 200 nm aer
modication with the copolymer which corresponded to the
TEM images.
Drug loading and release
DOX was chosen as a model anti-cancer drug to discuss the
drug loading and release behaviour of PRMS-FA/HMS/C18.
DOX was rst loaded into the hollow core of HMS/C18. The N2
adsorption/desorption isotherms of HMS, HMS/C18 and
DOX/HMS/C18 (Fig. S4†) indicated that the modication of
C18 and loading of DOX molecules had little inuence on the
pore structure of the HMS, which is signicant for drug
loading and release. The drug loading content was deter-
mined by measuring the DOX concentration before and aer
loading the HMS/C18. The theoretical loading content was set
at 5%, 10% and 50%, respectively and the DOX loading effi-
ciency was measured as 74.6, 73.4 and 66.8% accordingly
(Table 1). It was obvious that HMS had higher drug loading
content and efficiency than common mesoporous silica due to
its hollow core.
Evaluation of cytotoxicity and cell uptake
The sulforhodamine B (SRB) assay was used to assess the
cytotoxicity of the nanocarrier. Aer incubation with the
nanocarrier for 72 h, the cells displayed high cell viability
(>80%) as shown in Fig. 5. Even at a high concentration, the
cytotoxicity still remained at a low level, which is signicant for
the later experiment in vitro. The results above proved that the
polymer shell was biocompatible. These characteristics are
necessary for further biomedical application.
To investigate the properties of HMS/C18/PRMS-FA nano-
carriers in tumor cell targeting and cell imaging, the cellular
uptake of nanocarriers by FR(+) and FR(ꢁ) cells was measured.
The DOX-loaded HMS/C18 was coated with the copolymer
PRMS-FA via a self-assembly process to obtain DOX/HMS/C18/
PRMS-FA. And the results of DOX release from HMS/C18 and
HMS/C18/PRMS-FA indicated that the copolymer layers could
block the nanopores of HMS to cage the pre-loaded drugs
(Fig. S5†).
Table 1 Drug loading content and drug loading efficiency of HMS/C18
Theoretical drug loading
content (wt%)
Drug loading
Drug loading
efficiency (%)
contenta (wt%)
5
10
50
3.73
7.34
33.4
74.6
73.4
66.8
a
Fig.
5 In vitro cell viability of HMS/C18/PRMS-FA at different
Drug loading content for DOX was determined by UV–Vis
spectroscopy.
concentrations.
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nanocarriers aer 2 h. The changes in uorescence intensity
suggested that the targeting moiety offered by folic acid is effi-
cient at enhancing tumor cell targeting in vitro. The high uo-
rescence intensity of the nanocarriers ensures their application
can be utilized in uorescence imaging.
Observation of the FRET process in cancer cells by CLSM
We investigated use of the FRET property for monitoring the
light-controlled drug release from the nanocarriers by confocal
laser scanning microscopy (CLSM). We incubated the DOX-
loaded nanocarriers (500 mg mLꢁ1) with KB cells for 2 h in the
dark at rst. Then the cells were exposed to UV light (l ¼
365 nm, 0.2 W cmꢁ2) for time periods ranging from 0 to
20 minutes and the cells were examined by CLSM. As can be
seen from Fig. 7, a strong red emission excited at 510 nm was
visible in the KB cells without UV irradiation while the uo-
rescence intensity decreased gradually with as the UV irradia-
tion time increased. The mean uorescence intensity of cancer
cells is shown in Fig. S6.† We attributed the change in uo-
rescence intensity to the structure conversion of the light-
responsive copolymer in the nanocarriers, resulting in the FRET
process. Since the structure conversion would make the light-
responsive copolymer break away from the HMS nanoparticles,
triggering the drug release, the FRET process combined with
the light-controlled drug release has the ability to be a real-time
monitor.
Fig. 6 CLSM of FR (ꢁ) A549 cells and FR (+) KB cells incubated with
DOX/HMS/C18/PRMS-FA for 0.5 h and 2 h.
The in vitro cytotoxicity of DOX/HMS/C18/PRMS-FA under UV
irradiation was evaluated using the MTT assay. We incubated
KB cells in a culture medium with DOX/HMS/C18/PRMS-FA
under UV irradiation, and with HMS/C18/PRMS-FA, DOX/HMS/
C18/PRMS-FA, and a blank sample under UV irradiation as the
controls. Aer 3 h of incubation, the nanocarriers were removed
and the cells were incubated for another 24 h in the dark. As can
be seen in Fig. 8, treatment with UV light at a low power did not
result a signicant decrease in the cell viability. The cells still
kept high viability aer being treated with DOX/HMS/C18/
PRMS without UV irradiation, indicating no DOX was released
without irradiation. In contrast, DOX/HMS/C18/PRMS-FA under
UV irradiation within 20 minutes showed an obviously
enhanced cytotoxicity to the KB cells, which could be explained
by the light controlled drug release.
Fig. 7 CLSM observations of the FRET process.
Conclusion
In summary, we have prepared a core–shell nanocomposite as
nanocarrier and real-time monitor for light controlled drug
delivery via a facile strategy. HMS, an ideal base material with
good biocompatibility, was introduced into the nanocarrier to
enhance its drug loading capability and stability. By introducing
selective targeting folate groups to the copolymer, the nano-
Fig. 8 Inhibition of KB cell growth in the presence of excitation at 365
nm (20 min), drug carrier HMS/C18/PRMS-FA and DOX loaded drug
carrier without or with UV light exposures for varying durations after 24
h of incubation.
The confocal laser scanning microscopy (CLSM) images of cells carrier could target into FR(+) KB cells. The UV light (l ¼ 365 nm)
with DOX/HMS/C18/PRMS-FA solution at different incubation could make the amphiphilic copolymer shell become hydrophilic
times are shown in Fig. 6. Aer 0.5 h, there was no signicant and separate from the hydrophobic surface of the HMS nano-
difference in uorescence intensity between the two nano- particles, resulting in a signicant amount of drug release (>70%)
carriers. However, a signicant increase in uorescence inten- to inhibit the cancer cells. Moreover, a FRET process based on the
sity can be observed in KB cells incubated with FA-conjugated transformation of the copolymer occurred simultaneously with
1188 | J. Mater. Chem. B, 2014, 2, 1182–1189
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Journal of Materials Chemistry B
the drug release process which could provide an approach to
monitor the light-triggered drug release in real time. In vitro
studies conrmed the biocompatibility and the efficacy of this
nanocarrier for cancer cell targeting and inhibition, which
guarantees it a potential biomedical application.
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We gratefully acknowledge the nancial support provided by
National Natural Science Foundation of China (21176164,
21301125), Natural Science Foundation of Jiangsu Province
(BK2012625), Natural Science Foundation of the Jiangsu Higher
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