A multifunctional lipid that forms contrast-agent liposomes with dual-control release capabilities for precise MRI-guided drug delivery

Article abstract of DOI:10.1016/j.biomaterials.2019.119412

Monitoring of nanoparticle-based therapy in vivo and controlled drug release are urgently needed for the precise treatment of disease. We have synthesized a multifunctional Gd-DTPA-ONB (GDO) lipid by introducing the Gd-DTPA contrast agent moiety into an o-nitro-benzyl ester lipid. By design, liposomes formed from the GDO lipid combine MRI tracking ability and dual-trigger release capabilities with maximum sensitivity (because all lipids bear the cleavable moiety) without reducing the drug encapsulation rate. We first confirmed that both photo-treatment and pH-triggered hydrolysis are able to cleave the GDO lipid and lyse GDO liposomes. We then investigated the efficiency of drug release via the combined release processes for GDO liposomes loaded with doxorubicin (DOX). Relative to neutral pH, the release efficiency in acidic environment increased by 10.4% (at pH = 6.5) and 13.3% (at pH = 4.2). This pH-dependent release response is conducive to distinguishing pathological tissue such as tumors and endolysosomal compartments. The photo-induced release efficiency increases with illumination time as well as with distance of the pH from neutral. Photolysis increased the release efficiency by 13.8% at pH = 4.2, which is remarkable considering the already increased amount of drug release in the acidic environment. In addition, the relaxation time of GDO liposomes was 4.1 times that of clinical Gd-DTPA, with brighter T₁-weighted imaging in vitro and in vivo. Real-time MRI imaging and in vivo fluorescence experiments demonstrated tumor targeting and MRI guided release. Furthermore, significant tumor growth inhibition in a treatment experiment using DOX-loaded GDO liposomes clearly demonstrated the benefit of photo-treatment for efficacy: the tumor size in the photo-treatment group was 3.7 times smaller than in the control group. The present study thus highlights the benefit of the design idea of combining efficient imaging/guiding, targeting, and triggerable release functions in one lipid molecule for drug delivery applications.

Full text of DOI:10.1016/j.biomaterials.2019.119412

Journal Pre-proof
A multifunctional lipid that forms contrast-agent liposomes with dual-control release
capabilities for precise MRI-guided drug delivery
Chenyu Liu, Kai K. Ewert, Ning Wang, Youli Li, Cyrus R. Safinya, Weihong Qiao
PII:
S0142-9612(19)30511-3
DOI:
https://doi.org/10.1016/j.biomaterials.2019.119412
JBMT 119412
Reference:
To appear in: Biomaterials
Received Date: 9 April 2019
Revised Date: 18 July 2019
Accepted Date: 3 August 2019
Please cite this article as: Liu C, Ewert KK, Wang N, Li Y, Safinya CR, Qiao W, A multifunctional lipid
that forms contrast-agent liposomes with dual-control release capabilities for precise MRI-guided drug
delivery, Biomaterials (2019), doi: https://doi.org/10.1016/j.biomaterials.2019.119412.
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A multifunctional lipid that forms contrast-agent liposomes with
dual-control release capabilities for precise MRI-guided drug
delivery
Chenyu Liu1,2, Kai K. Ewert2, Ning Wang1, Youli Li3, Cyrus R. Safinya2*, Weihong Qiao1*
1
State Key Laboratory of Fine Chemicals, School of Chemical Engineering, Dalian University of
Technology, Dalian, 116024, P. R. China
2
Materials, Physics, and Molecular, Cellular & Developmental Biology Departments, University of
California at Santa Barbara, Santa Barbara, California 93106
3 Materials Research Laboratory, University of California at Santa Barbara, Santa Barbara, California
93106
*Corresponding authors contact information:
qiaoweihong@dlut.edu.cn
Safinya@mrl.ucsb.edu
TOC graphic
1

Abstract
Monitoring of nanoparticle-based therapy in vivo and controlled drug release are urgently
needed for the precise treatment of disease. We have synthesized a multifunctional
Gd-DTPA-ONB (GDO) lipid by introducing the Gd-DTPA contrast agent moiety into an
o-nitro-benzyl ester lipid. By design, liposomes formed from the GDO lipid combine MRI
tracking ability and dual-trigger release capabilities with maximum sensitivity (because all
lipids bear the cleavable moiety) without reducing the drug encapsulation rate. We first
confirmed that both photo-treatment and pH-triggered hydrolysis are able to cleave the GDO
lipid and lyse GDO liposomes. We then investigated the efficiency of drug release via the
combined release processes for GDO liposomes loaded with doxorubicin (DOX). Relative to
neutral pH, the release efficiency in acidic environment increased by 10.4% (at pH = 6.5) and
13.3% (at pH = 4.2). This pH-dependent release response is conducive to distinguishing
pathological tissue such as tumors and endolysosomal compartments. The photo-induced
release efficiency increases with illumination time as well as with distance of the pH from
neutral. Photolysis increased the release efficiency by 13.8% at pH=4.2, which is remarkable
considering the already increased amount of drug release in the acidic environment. In
addition, the relaxation time of GDO liposomes was 4.1 times that of clinical Gd-DTPA, with
brighter T1-weighted imaging in vitro and in vivo. Real-time MRI imaging and in vivo
fluorescence experiments demonstrated tumor targeting and MRI guided release. Furthermore,
significant tumor growth inhibition in a treatment experiment using DOX-loaded GDO
liposomes clearly demonstrated the benefit of photo-treatment for efficacy: the tumor size in
the photo-treatment group was 3.7 times smaller than in the control group. The present study
thus highlights the benefit of the design idea of combining efficient imaging/guiding,
targeting, and triggerable release functions in one lipid molecule for drug delivery
applications.
Keywords
Cancer therapy, Dual-control release, MRI-guided drug delivery, Nanoparticles,
Contrast-agent liposomes
2

1. Introduction
Nanoscale liposome delivery systems have been widely explored for biomedical
applications,[1] including for the diagnosis of tumors,[2-3] tumor therapy,[4-5] treatment of
inflammation,[6] and central nervous system medication[7-8] because they possess desirable
properties such as targeting, high encapsulation rates, sustained release ability and
biocompatibility. The applications of liposomes include both clinical trials and approved
drugs.[9-10] Nanocarriers allowing for precisely adjustable release are highly desirable to
improve on current therapeutics because they will allow minimizing the side effects (on
healthy cells and tissues) of disease treatments. To meet this demand of simultaneously
reducing the toxicity of a drug and improving its efficacy, numerous approaches to trigger
release from drug-loaded nanoliposomes at the precise site of a lesion or disease have been
derived; release can be controlled, for example, by ultrasound,[11] pH,[12-13] light,[14]
singlet oxygen,[15-16] heat,[17] or redox potential.[18] However, the complexity of the
environment in vivo[19-20] can interfere with the desired release effect, and not all of these
approaches have met with therapeutic success.[21] Here, we present a strategy to address the
challenge of selective delivery by developing a targeting nanoscale liposome with triggered
release of the loaded drug and real-time tracking capabilities to monitor its transport
trajectory in vivo. Thus, when the tracking capabilities show that the liposomes have been
delivered to the lesion, the external trigger is employed to rapidly disrupt the carrier and
release the drug, reducing unspecific, unwanted toxicity and improving therapeutic efficacy
at the desired site.
The strong anaerobic metabolism of tumor tissue and inflammatory lesions renders the
extracellular environment weakly acidic (pH 6.0–6.8). Inside cells, the processing of
endosomes into lysosomes is also accompanied by a drop in pH inside these organelles (to
pH 4.0–6.0). Thus, pH-sensitive lipids have received widespread attention. Most of the
reported pH-responsive liposomes[22-24] are designed such that low pH disrupts the carrier
and releases the loaded drug or gene. However, such single-response controlled-release
3

systems leave much room for improvement as they only increase the release efficiency by
around 20% [25-27]. Moreover, precision medicine may require the ability to modify the
release properties to accommodate the degree of inflammation or specific tumor properties.
Single-response drug release systems cannot meet these demands.[28-29]
While the efficiency of light-triggered release is limited,[30-32] this method has the
advantages of simple triggering conditions, high local precision of triggering, and minimal
damage to healthy tissue. Therefore, combining a light-triggered cleavable moiety together
with a pH-sensitive cleavage site will allow an added level of externally controlled
modification of drug release. The o-nitrobenzyl ester bond is easily cleaved by exposing it to
UV light, and its photocleavage proceeds via Norrish type II intramolecular rearrangement
without the need for an additional reactant or solvent.[33] Furthermore, it has been
demonstrated that o-nitrobenzyl micelles can be photo-disaggregated with near-infrared (NIR)
light by absorbing two NIR photons.[34-37] This greatly expands the utility of the
o-nitrobenzyl group because tissue is largely transparent to NIR light.[38-40] It is also
possible to encapsulate upconversion nanoparticles (UCNPs, anti-Stokes materials)[41-43] to
expand the applications of liposomes based on an o-nitrobenzyl lipid. An additional
advantage of the o-nitrobenzyl ester and related bonds is that it can combine multiple
functions in a simple structure[44-45] (including pH-sensitivity, courtesy of the ester bond).
Thus, the o-nitrobenzyl ester moiety can be expected to promote pH-triggered initial drug
release which can then be manually and precisely enhanced by tailored phototherapy
conditions to locally dissociate the liposomes.
Despite decades of research and progress, cancer remains a major challenge for medicine,
accounting for 22.5% of total mortality[46]. While chemotherapy is a relatively mild
treatment compared to surgery and radiation therapy, the targeted delivery of anticancer
drugs is a promising strategy to further reduce side effects. While not containing specific
targeting moieties, our lipid-based drug carrier is designed to exploit the "enhanced
4

permeability and retention” (EPR) effect discovered by Maeda and Matsumura.[47] The EPR
effect applies to almost all rapidly growing solid tumors[48] and has been observed for
biocompatible macromolecular compound above 40 kDa as well as nanoparticles above a
certain size. Self-assembled nanoparticles such as micelles and liposomes containing
anticancer drugs are excellent candidates for targeting via the EPR effect,[49-50] with
measured concentrations of the drug 10–30 times higher in the tumor than in the blood.
A key design feature of our liposomal delivery system is the ability to monitor the transport
path of the liposomes in real time in vivo,[51] to more accurately determine the timing of the
phototherapy and thereby control drug release and reduce toxicity. Several approaches to
achieve this monitoring may be considered. Fluorescence has been used to monitor drug
release based on changes in Förster Resonance Energy Transfer signal in nanoparticles.
However, the application of this method in vivo is limited by tissue penetration of the exciting
and emitted light.[52-53] Isotopic labelling, i.e., marking the nanocarrier or drug with
radioisotopes, enables monitoring of the drug delivery system, but the undesirable and
detrimental effects of radiation limit the application of this method.[54-55] Magnetic
resonance imaging (MRI), by virtue of excellent tissue penetration, lack of hazardous
radiation, and a high degree of imaging contrast, has the most potential for in vivo monitoring
of drug delivery systems.
In recent years, a number of MRI-guided controlled-release liposomes have been reported,
which can largely be classified into two categories. One category comprises liposomes that
encapsulate a small amount of contrast agent, such as manganese-based compounds,[56-57]
gadolinium-chelating materials,[58-61] or superparamagnetic iron oxide.[62-63] The
disadvantage of this approach is that it reduces the drug encapsulation efficiency.[64-66] The
second approach is mixing triggerable lipids with lipids that act as MRI contrast
agents.[67-68] This approach effectively retains the high drug encapsulation efficiency but
weakens the efficiency of release because the triggerable lipid is diluted.[69-70] Because
5

both these approaches reduce efficacy, a need for improvement of MRI-guided
controlled-release liposomes remains.[71]
Here we present a new approach of combining tracing and dual-trigger controlled-release
capabilities in a single lipid molecule (Gd-DTPA-ONB lipid, abbreviated as GDO lipid) to
create a drug carrier for precision medicine. The molecule contains two lipid tails (to allow
liposome formation), an o-nitrobenzyl ester (for dual trigger response), and a linked MRI
contrast agent (Gd-DTPA, for MRI-surveillance of the movement of the drug-loaded
liposome). The contrast agent allows using imaging to ensure precise manual control of drug
release to fit the severity of the condition. As an additional benefit, attaching Gd-DTPA to a
hydrophobic moiety is expected to significantly increase the contrast of the MRI agent
according to the Solomon-Bloembergen-Morgan theory (SBM theory),[72] by raising the T1
MRI contrast and lateral relaxation rate (R1).
As shown in Scheme 1a, our design approach results in the surface of the assembled liposome
being completely covered by the headgroup containing the MRI contrast agent (Gd-DTPA),
which strongly enhances the MR signal of the liposome without compromising the
encapsulation ability. In addition, having multiple trigger responses at the same site makes
the liposomes more sensitive than other complex controlled-release liposomes. Initially, GDO
liposomes release part of the loaded drug due to triggering by the pH. After that, photo
triggering can release additional drug in a controlled manner.
Scheme 1b schematically shows the application of this approach in cancer therapy. After
intravenous injection of drug-loaded GDO liposomes in a tumor mouse model, the transport
pathway of the liposomes is detected by MRI. When the drug-loaded GDO liposomes display
enrichment at the tumor site by the EPR effect with significant brightening in MRI imaging,
the tumor sites are irradiated with light to facilitate dissociation of the liposomes beyond the
initial pH-triggered response.
6

We
synthesized
the
GDO
lipid
starting
from
didodecylamine
and
4-(bromomethyl)-3-nitro-benzoic acid, with partial hydrolysis of the dianhydride of DTPA
yielding the key precursor of the Gd-DTPA moiety. After confirming the sensitivity of both
the GDO lipid and GDO liposomes to pH- and photo-triggered cleavage, we successfully
encapsulated doxorubicin (DOX) as a model hydrophobic drug. After comparing the
cytotoxicity of empty and DOX-loaded GDO liposomes, we investigated the photo- and
pH-triggered release of DOX from these liposomes. We further assessed the utility of MRI to
track the pathway and tumor targeting ability of GDO liposomes, using fluorescence imaging
(with DOX-loaded GDO liposomes) to verify the obtained results. Encouraged by the high
MRL contrast, tumor targeting ability and controlled release properties of the GDO liposomes,
we investigated their therapeutic utility in a xenograft mouse model of breast cancer. We used
UV light to trigger release from the GDO liposomes because the tumor model was close to
the epidermis and found that tumor growth was inhibited the most in animals treated with
DOX-loaded GDO liposomes in combination with phototherapy.
a
pH
Gd-DTPA-ONB lipid
hν
DOX/GDO Liposome
Initial cleavage:
Photo-triggering to adjust
the degree of cleavage
EPR; High MRI contrast
pH-triggered release
DOX·HCl
7

b
pH
pH release
Photodynamic release
MRI
Tracing
Tumor
Tail vein injection
Targeting enriched at tumor
Illumination ray
Mitochondrion
Golgi apparatus
Released DOX
Blood vessel
Nucleus
4-3
Capillaries
Tumor Cell
4-2
4-1
Lysosome
2-3
1-1
3-3
2-2
3-2
Endosome
Centriole
2-1
3-1
DOX/GDO
DOX/GDO
Cleaved DOX/GDO
Red blood cell
Enriched DOX/GDO liposomes (EPR)
with MRI imaging
pH triggered & manually photo-adjusted release
Tumor tissue
Tumor Cell
Scheme 1. Schematic depiction of the use of GDO liposomes for drug delivery. The drug-loaded GDO
liposomes combine the functions of contrast agent for MRI-guided therapy, EPR effect-based targeting,
and multistage drug release. (a) Formation and progressive cleavage of drug-loaded GDO liposomes. GDO
lipids and the encapsulated drug (doxorubicin hydrochloride, DOX · HCl) self-assemble to form
nanometer-size drug-loaded liposomes (DOX/GDO liposomes). Since the entire surface of the GDO
liposomes is covered by the Gd-DTPA headgroup of the GDO lipid, these passively targeted liposomes are
a highly efficient MRI contrast agent, allowing in vivo tracking. Some of the o-nitrobenzyl ester linkages
in the GDO lipids will be cleaved by a low-pH environment, resulting in some initial release of drug.
Subsequently, the extent of drug release can be adjusted by controlling the conditions of irradiation for
photo-triggered release. (b) Overview of the use of GDO liposomes for cancer therapy. After intravenous
injection of drug-loaded GDO liposomes in a mouse tumor model, the transport pathway of the
high-contrast liposomes is followed by MRI. As the liposomes enrich at the tumor site by the EPR effect,
the acidic extracellular environment of the tumor stimulates an initial release of DOX which transfers into
the cells as free drug (1-1). Meanwhile, most DOX/GDO liposomes enter the cell by endocytosis (2-1). As
8

the endosomes (2-1) are processed into lysosomes their interior becomes highly acidic, leading to GDO
cleavage and the release of free DOX (2-2), followed by transfer of the free drug into the nucleus (2-3). At
the same time, because DOX/GDO liposome accumulation at the tumor site is detected with MRI,
photo-triggered release is initiated by irradiation of the tumor site. This releases additional free DOX from
extracellular DOX/GDO liposomes (3-1) to enter the cell by osmotic pressure (3-2) and transfer to the
nucleus (3-3). In addition, release of DOX from endocytosed DOX/GDO liposomes (4-1) is accelerated by
the synergism of light and lysosomal low pH (4-2), and the released DOX transfers to the cell’s nucleus
(4-3).
2. Experimental Section
2.1 Photocleavage of Gd-DTPA-ONB
A small amount of Gd-DTPA-ONB lipid was dissolved in a closed quartz sample cell
(chloroform/methanol = 1:2, v/v). Samples of this solution were irradiated with UV light (365
nm, 60 W) for 0, 5, 10, 30, 60 min. After irradiation, TLC was performed to assess the degree
of photocleavage, and MS was used to detect the compounds in the sample irradiated for 60
min.
2.2 pH-sensitivity of Gd-DTPA-ONB
Equal amounts of Gd-DTPA-ONB lipid were dissolved in equal volumes of phosphate buffer
solution with different pH (0.01 M, pH 4.0–13.0). After reacting at room temperature for 1 h,
the
water
in
the
sample
was
removed
by
lyophilization.
TLC
(chloroform/methanol/water/acetic acid = 65:25:4:1, v/v) was used to assess cleavage of the
molecules, and MS was used to analyze the samples. The cleavage product observed at the
solvent front in TLC was collected and separated in a preparative assay by chromatography
on silica gel, using a step-wise gradient of chloroform/methanol mixtures (30:0 to 6:1, v/v).
The resulting colorless oil was characterized by TLC (chloroform/methanol = 6:1 v/v; Rf =
0.69), MS, and NMR.
2.3 Preparation of empty Gd-DTPA-ONB liposomes
A total of 60.7 mg of Gd-DTPA-ONB lipid was dissolved in 5 mL of chloroform/methanol
9

(1:2, v/v) to prepare a 10 mM stock solution. An aliquot of 500 ꢀL of the stock solution in a
glass vial was rotary evaporated at room temperature under nitrogen to obtain a thin lipid film.
After drying in a vacuum for 16 h, a total of 1 mL of 0.01 M PBS was added to the vial and
the mixture was incubated at 37 °C for 24 h and then sonicated with an ultrasonic signal
generator for 10 min (ultrasound 2 s, intermittent 5 s, power 45%). This process was repeated
(using varying amounts of PBS) to prepare a range of GDO liposome sample concentrations:
5.0 mM, 2.5 mM, 1.25 mM, 1.00 mM, 0.625 mM, 0.315 mM, 0.156 mM, 0.078 mM, and
0.039 mM.
2.4 Characterization and photolysis behavior of empty GDO liposomes
At 12 h after preparation, aliquots of a 1 mM suspension of empty GDO liposomes in PBS
was sealed in a quartz tube and irradiated for 0, 5, 10, 20, 30, 40, 60, 90, and 120 min (365
nm at 60 W). The hydrodynamic diameter and size distribution of the empty GDO liposomes
in solution was then measured using a Malvern Zetasizer Nano ZS. To directly observe the
morphological changes of GDO liposomes in real time, differential interference contrast
microscopy was used to monitor the samples at different illumination times. Transmission
electron microscopy (TEM) samples were prepared by placing 30 ꢀL of sample solution on
300-mesh copper grids after light exposure for 0 min and 120 min. After drying at room
temperature, the copper mesh was immersed in a droplet of phosphotungstic acid (2%) for
negative staining for 3 min, removed, dried, and imaged using a JEOL-JEM2000EX electron
microscope.
2.5 Self-assembly of DOX/GDO liposomes and encapsulation efficiency at varied
lipid/drug ratio
Aliquots of the stock solution of Gd-DTPA-ONB were evaporated under a stream of nitrogen
at room temperature to obtain a thin lipid film. After drying in a vacuum overnight, varying
amounts of a stock solution of 500 mg/L DOX·HCl in PBS (0.01 M; pH 7.4) were added to
obtain samples at different lipid/drug ratios. The samples were incubated at 37 °C for 22 h
and then sonicated for 10 min (ultrasound 2 s, intermittent 5 s, power 45%). The resulting
10

suspension was transferred to a dialysis bag (MW cutoff=3500) and immersed in 500 mL of
PBS buffer (0.01 M; pH 7.4) with stirring at 37 °C. Dialysis was performed for 4 h while
exchanging the PBS buffer every 0.5 h. The liposome suspension was then removed and
dried by lyophilization, and the resulting solid was dissolved in 5 mL of DMF to measure the
drug encapsulation efficiency by referring to the DOX fluorescence emission standard curve.
2.6 Characterization and photolability of DOX/GDO liposomes
To assess liposome stability, samples of sonicated DOX/GDO liposomes were incubated in
the dark for 5 days, and particle size was measured daily using a Malvern Zetasizer Nano ZS.
To assess the effect of light irradiation on liposome stability, samples of DOX/GDO liposome
solution were irradiated for 0, 5, 10, 20, 30, 40, 60, 90, and 120 min (365 nm at 60 W) in
closed quartz tubes. The hydrodynamic diameter and size distribution of the DOX/GDO
liposomes were measured using a Malvern Zetasizer Nano ZS. In addition, TEM
(JEOL-JEM2000EX and HT7700 EXALENS electron microscopes) was used to assess the
morphological changes in the drug-loaded liposomes. To this end, we imaged samples before
illumination, and the supernatant and sediment suspension of the samples illuminated for 120
min.
2.7 Light-triggered drug release from DOX/GDO liposomes
To assess the photo-triggered release of DOX from drug-loaded liposomes, free DOX was
first removed from DOX/GDO liposomes by dialysis. The drug-loaded liposome suspension
was then transferred quartz tubes, and 1 mL each was irradiated with UV light (365 nm at 60
W) for a different time (0, 10, 20, 30, 40, 60, 90, and 120 min). The solutions as well as an
amount of free DOX equal to the liposome-encapsulated amount free were placed in fresh
dialysis bags (MW=3500) and then immediately immersed in 80 mL of PBS buffer at 37 °C
with stirring. Aliquots of 3 mL of the PBS buffer were withdrawn at appropriate intervals and
the fluorescence emission intensity at 596 nm under excitation of 478 nm was measured by
fluorescence spectrometry to calculate the DOX concentration and obtain the cumulative
release curve as follows:
11

ꢂ
ꢃ
ꢀ ꢁ
Cumulative release of DOX % =
×ꢅꢆꢆ%
ꢂ
ꢄ
Here, Mt is the amount of DOX released after time t, and M0 is the amount of DOX initially
encapsulated in the liposome.
The release efficiency was calculated as
ꢇꢈ
ꢃ
ꢀ ꢁ
Release efficiency % =
×ꢅꢆꢆ%
ꢇꢈ
ꢄ
Here, CRt is the cumulative release of DOX for a given time or sample, and CR0 is the total
cumulative release of DOX from the free DOX sample at same pH value.
2.8 Measurement of pH- and light-triggered DOX release from DOX/GDO liposomes
To simulate pH- and light-triggered release drug from DOX/GDO liposomes in vivo, a 1 mM
DOX/GDO liposome suspension was prepared at neutral pH, divided into 3 equal parts and
irradiated for three different exposure times (0, 5, and 20 min). Then each irradiated group
was placed in quadruplicate into dialysis tubes and immersed in 80 mL phosphate buffer at
various relevant pH values (pH= 4.2, 6.5, 7.4, 8.5) at 37 °C with stirring. In addition to free
DOX, 1 mM DOX/DOPC liposomes (stable to UV and pH) were used as a control.
Cumulative DOX release and release efficiency were measured and calculated as described
above.
2.9 Cytotoxicity of empty GDO liposomes
MCF-7, Hela7702, HepG2 cells and Raw 264.7 macrophages in the logarithmic growth phase
were seeded in 96-well plates at an approximate cell density of 105 cells/mL. The 96-well
plates were placed in an incubator at 37 °C (5% CO2 atmosphere). After the cells were
completely attached and had reached about 80% confluence, 80 ꢀL of fresh medium
containing 20 ꢀL of GDO liposome supension (at 5 mM, 2.5 mM, 1.25 mM, 0.625 mM,
0.315 mM, 0.156 mM, 0.078 mM, and 0.039 mM) was added. After incubation in the
incubator for 24 h, the medium was aspirated and the cells were washed twice with fresh PBS
buffer. A total of 150 ꢀL of a solution of MTT (0.5 mg/mL) was added and incubation was
continued for 4 h. The MTT reagent solution was then carefully removed to obtain violet
crystals which were dissolved in 150 ꢀL of DMSO. The 96-well plate was then placed in a
12

microplate reader, shaken for 10 min, and the absorbance at 492 nm was measured. Cell
survival was obtained as Asample/Acontrol, where Acontrol is the absorbance measured in a well
that contained untreated cells.
2.10 Cytotoxicity of DOX-loaded liposomes
To MCF-7, Hela7702, HepG2 cells and Raw 264.7 macrophages, seeded as described above
(2.9) and at 80% confluency, 80 ꢀL of medium containing 20 ꢀL of a 1 mM suspension of
DOX/GDO liposomes were added. Treated cells, as well as untreated control cells, were
irradiated with UV light (365 nm at 60 W) for 0, 5, 10 and 20 min, respectively. For
comparison, another set of cells was treated with a solution of free DOX equal in amount and
DOX concentration to the liposome suspension. The cells were then incubated for 24 h and
cell viability measured using the MTT assay as described above.
2.11 Magnetic relaxivity performance of GDO in vitro and in vivo
(All animal experiments comply with the ARRIVE guidelines and be carried out in
accordance with the National Institutes of Health guide for the care and use of Laboratory
animals)
Different concentrations of GDO and Gd-equimolar amounts of Gd-DTPA were used for
comparison. The sample relaxation time and in vitro imaging were measured using a
MesoMR23-060H-I magnetic resonance imager. The relaxation efficiency was obtained by
fitting the value of 1/T1 to the gadolinium concentration. To further investigate the nuclear
magnetic resonance behavior of the empty GDO liposomes in vivo, pathogen-free Dutch mice
(~26 g) were anesthetized with a subcutaneous injection of 150 ꢀL of 7% chloral hydrate. A
total of 200 ꢀL of a 1 mM GDO liposome suspension or a Gd-equivalent amount of
Gd-DTPA solution was injected via the tail vein. The T1-weighted image was recorded 0, 15,
30, 45, 60, 120, 180, and 240 min after injection. Through the coronal and sagittal planes of
the intuitive MRI, the changes of brightness in the lungs, heart, liver, and kidneys of mice
were observed. The MesoQMR23-060H-I magnetic resonance imager was used at a
resonance frequency of 21.245 MHz, a magnet strength of 0.5 T, a coil diameter of 40 mm,
13

and a magnet temperature of 32 ± 0.01 °C. T1-weighted image acquisition parameters: FOV
Read=100 mm, FOV Phase=100 mm, TR=500 ms, TE=20.000 ms.
2.12 Cancer treatment in a nude mouse model of human breast cancer
Human breast cancer cells (MCF-7; 5×107 cells) were directly injected into the right anterior
leg of 5–6 week-old specific pathogen-free nude mice. After 3 weeks, tumor size reached
approximately 90 mm3.[73] The mice were randomly divided into 5 groups. The first group
of mice received PBS (control group); the second group was injected with empty GDO
liposomes (“GDOL group”); the third group was injected with DOX solution (“DOX group”);
the fourth group was injected with DOX/GDO liposomes (“DOX/GDO group”); the fifth
group was injected with DOX/GDO liposomes and received light treatment (“DOX/GDO-UV
group”). Injections were performed every 2 days via the tail vein, and the DOX-equivalent
was 2.5 mg/kg for each group receiving DOX (DOX, DOX/GDO, and DOX/GDO-UV
groups). For the DOX/GDO-UV group, only the tumor site was intermittently irradiated for 8
min using UV light (365 nm at 60 W, irradiated for four rounds of 2 min on, 3 min off) at 30
min after injection of DOX/GDO liposomes. The body weight of the mice was recorded, and
the long and short diameters of the tumor were measured with a vernier caliper to assess the
effect of the treatments. The volume of the tumors was calculated as V = a×b2/2, with a the
long diameter of the tumor and b the short diameter of the tumor.[74] After 16 d of
treatment,[75] the mice were sacrificed and dissected. Heart, liver, spleen, lung, kidney and
tumor tissues were subjected to hematoxylin-eosin staining to evaluate systemic toxicity. The
morphology of the tissues was observed under an inverted microscope, and the distribution of
DOX in the tumor tissues was observed by fluorescence microscopy.
2.13 In vivo MRI-tracking studies
To test MRI tracking and targeted aggregation capabilities, nude mice with a tumor volume of
approximately 1800 mm3 (one per treatment protocol) received a single treatment as
described above for the Gd-DTPA, DOX/GDO and DOX/GDO-UV groups. An
MesoQMR23-060H-I magnetic resonance imager was used to track the real-time transport
14

path of Gd-DTPA and DOX/GDO liposomes in vivo, and the contrast imaging brightness of
the tumor sites in the nude mice was observed. T1-weighted imaging was performed 0, 15, 30,
45, 60, 120, and 240 min after injection, at a resonance frequency of 23.313 MHz, a magnet
strength of 0.5 T, a coil diameter of 40 mm, and a magnet temperature of 32 ± 0.01 °C
(T1-weighted imaging parameters: FOV Read=100 mm, FOV Phase=100 mm, TR=360 ms,
TE=18.125 ms).
2.14 In vivo fluorescence imaging studies
Fluorescence in vivo imaging studies of the DOX, DOX/GDO and DOX/GDO-UV groups
were performed as a secondary method of verifying the transport path and the tumor targeting
capabilities of GDO liposomes. Mice were treated as described for the tumor treatment
protocol, and then real time imaging using a Berthold LB983NC100 in vivo imaging system
was performed to monitor the fluorescence from DOX.[76-80] At the selected time point (4h
or 24 h), the mice were sacrificed, and tumors and major organs were collected, washed and
imaged to determine the biodistribution of DOX.
3. Results and Discussion
3.1. Design and synthesis of the high-MRI contrast, multistage-responsive
Gd-DTPA-ONB lipid
To enable precision therapy, preventing indiscriminate release in healthy tissues and thereby
reducing toxicity and side effects in particular, we sought to combine MRI contrast capability
for real-time monitoring, targeting, and tunable controlled-release functions into the design of
the GDO lipid. Integrating the Gd-DTPA MRI contrast agent moiety into the lipid structure
by chemical synthesis, rather than merely encapsulating a contrast agent or adding a
MRI-contrast lipid is one of the highlights of the design, as it promotes MR contrast without
reducing the controlled-release or drug encapsulation capabilities of the liposomes.
Additionally, the increase in the molecular weight of the MRI contrast agent helps to increase
relaxation efficiency according to the SBM theory.[72] The two long hydrophobic carbon
chains of GDO allow it to self-assemble and form liposomes that can encapsulate drugs for
15

delivery. This also serves to promote targeting because the typical size of liposomes formed
by small amphiphilic molecules with two hydrophobic chains is between 200 nm and 400 nm,
resulting in passive targeting via the EPR effect. To achieve a multistage release response that
can accommodate differences in lesion severity, our design incorporates automatic release
with pH as the trigger and adjustable release with light as the trigger. Light-triggered release
can be tuned in a variety of ways, such as changing the illumination time, light intensity, or
wavelength.[81-83] Utilizing an o-nitrobenzyl ester bond as the triggering moiety to achieve
both responses is another highlight of the design because it combines two functions in a very
simple structure.
The detailed synthesis route for GDO is shown in Scheme 2. Partial hydrolysis of
diethylenetriaminepentaacetic acid dianhydride (DTPAA) is the first key step of the synthesis:
converting as much of the dianhydride to the monoanhydride as possible facilitates
purification at a later stage because it limits the formation of various similar, polar
side-products. To begin building the hydrophobic part, didodecylamine was coupled with
4-(bromomethyl)-3-nitrobenzoic acid via amidation of the acid chloride. The triggerable
o-nitrobenzyl ester bond was then introduced by nucleophilic substitution of the benzyl
bromide moiety with the carboxyl group of Boc-Asp(O-allyl)-OH. Removal of the
Boc-protecting group from the resulting hydrophobic compound using trifluoroacetic acid
liberated an amine group which was then reacted with the above-mentioned “monoanhydride”
of DTPA to synthesize DTPA-ONB. Finally, DTPA-ONB was chelated with Gd(III) to form
the desired paramagnetic GDO lipid.
16

HOOC
N
O
HOOC
N
Acetic anhydride
Pyridine
H2O
HOOC
N
N
O
O
N
O
HOOC
HOOC
N
COOH
COOH
N
N
N
COOH
O
O
O
COOH
O
1
O
2
HN
NO2
Br
NO2
Br
NO2
SOCl2
Br
N
Et3N
O
Cl
3
4
COOH
O
O
N
O
O
DBU
O
COOH
H
O
O
NO2
O
H
O
TFA
O
N
O
O
N
O
2
NO2
O
O
O
O
6
5
O
HOOC
O
N
N
COOH
O
N
COOH
O
O
N
O
O
NO2
O
NO2
O
.
GdCl3 6H2O
O
O
O
N
N
O
HN
O
N
H
O
O
O
H
O
N
COOH
Gd
O
O
N
O
O
N
HOOC
8
7
O
N
COOH
COOH
O
N
Scheme 2. Synthesis of the GDO lipid; DBU: 1,8-Diazabicyclo[5.4.0] undec-7-ene; Et3N:
triethylamine; TFA: trifluoroacetic acid;
3.2 Characterization of the pH- and photo-responsiveness of the GDO lipid
Before investigating triggered release from self-assembled vesicles of GDO, we studied the
cleavage of the lipid by low pH and UV light at the molecular level. We used thin layer
chromatography (TLC) supplemented with MS and NMR to infer the dissociation mechanism.
To study the pH-responsiveness, GDO was dissolved in phosphate buffers of varied pH. TLC
showed (see Figure S12 in the Supporting Information) that the GDO lipid is stable at
physiological pH (pH = 7.4). However, at a pH lower than 7.0 or higher than 7.4, the
17

molecule was cleaved. The further the pH was from neutral/physiological pH (both higher or
lower), the greater was the observed extent of cleavage.[84] Mass spectrometry showed
(Figure S13 in the Supporting Information) that GDO hydrolyzed into two new compounds
(Scheme 3a) at pH=6.5, resulting in peaks at m/z=1216.83 (GDO (8)), m/z=716.55 ([M-H]– of
the gadolinium-chelated carboxylic acid derivative (11) and m/z=567.87 ([M+Cl]– of
N,N-didodecyl-4-(hydroxylmethyl)-3-nitrobenzamide (9; structure confirmed by MS and
NMR after the compound was separated by column chromatography (see Figures S14, S15 in
the Supporting Information)). We thus confirmed that GDO is a pH-sensitive lipid and that
the acid-driven cleavage proceeds as shown in Scheme 3a.
Next, we investigated the photo-triggered cleavage of GDO by irradiating samples with UV
light (365 nm, 60 W) and varying the illumination time. TLC analysis (see Figure S16 in the
Supporting Information) revealed that cleavage occurred already after 5 min of UV
irradiation and that the degree of cleavage increased with illumination time. Cleavage was
essentially complete after 120 min of irradiation as indicated by TLC and MS (see Figure S17
in the Supporting Information). The observed fragment peak at m/z=538.150 may be assigned
to N,N-di-dodecyl-4-formyl-3-nitrosobenzamide [M+Na]+, consistent with the photolysis
mechanism shown in Scheme 3b.[85-86]
The results described above demonstrate that GDO is susceptible to cleavage by two
triggering stimuli. In addition, the products of the light-induced cleavage may serve
additional anticancer purposes at the tumor sites beyond the dissociation of liposomes: the
fragment bearing a nitroso group is likely cytotoxic.[87]
18

(9)
Low or high
pH
+
[M+H]+
532.41
538.150
[M+Na]+
(11)
UV
+
(8)
(10)
H
O
O
H
O
O
NO2
O
O
O
H+
O
O
O
NO2
O
O
O
O
Gd
a
OH
O
+
.
N
O
N
N
N
N
HOOC
H
O
N
O
O
N
O
H
O
N
Gd
O
O
N
O
O
O
O
N
U
V
.. HO
.
O
.
O
O
.
O
O
OH
.
N
N
N
O
O
O
O
.
H
O
O
O
O
R1
O
R1
H
R1
O
O
O
R1=
R2=
O
R2
R2
R2
O
N
O
O
O
N
O
Gd
N
N
N
b
H
O
O
O
H
O
O
O
O
O
NO
H
O
O
O
O
N
O
N
O
O
O
N
Gd
H
+
N
O
HOOC
N
R1
N
O
R2
O
Scheme 3. (Top) The GDO lipid can be cleaved by two different triggers: pH and light. The
two triggers yield different lipid fragments. (Bottom) Mechanisms for the acid-catalysed
hydrolysis (a) and photolysis (b) of GDO.
3.3 The effect of UV irradiation on the morphology of GDO liposomes
After confirming that the GDO lipid is susceptible to photolysis, we set out to investigate
whether this lipid cleavage can initiate cargo release from drug-loaded liposomes. To this end,
we first studied the induced changes in morphology and size of empty GDO liposomes by
irradiating a 1 mM solution of empty GDO liposomes (prepared by hydration of a thin film)
with UV light for different times. Figure 1 compares the change in the particle size for an
unexposed sample (Figure 1a, with a slow growth from 190 nm to 310 nm over 5 days) and
the UV-irradiated group (Figure 1b, with a large and rapid size change). The diameter of the
liposomes in the UV-irradiated group initially was about 190 nm, with a relatively narrow
size distribution. With increasing exposure time, both the liposome size and the width of the
19

size distribution progressively increased. After 120 min of irradiation, the size of the
liposomes had increased nearly 10 times to 1300 nm. These observations are consistent with
real-time differential interference contrast microscopy data (see Figure S18 in the Supporting
Information) and the transmission electron microscopy (TEM) data described below. As
shown in Figure 1c, TEM reveals empty GDO liposomes to be regular bilayer spheres with a
size around 190 nm prior to UV irradiation. After 120 min of irradiation, the size of the
liposomes had increased to around 1100 nm (Figure 1d). The combined data demonstrates
that the vesicle morphology can be gradually modified by irradiation, suggesting that
photo-treatment could be a trigger for accelerating lysis of the liposomes and therefore drug
release. Compared to other work,[88-89] our GDO liposomes appear to be more responsive
to UV-irradiation. The likely reason for this is that the only constituent of the GDO liposomes
is the light-sensitive GDO molecule, meaning that there is no dilution of the o-nitrobenzyl
ester lipid in the membrane and the light sensitivity is therefore at its maximum possible level.
Based on the experimental data on size and morphology, we speculate that the mechanism
leading to the expansion of the liposome diameter is a process of dissociation and
reconstruction of liposomes, shown schematically in Figure 1e. UV light cleaves the GDO
lipid into ONB lipid and the Gd-DTPA derivative shown in Scheme 3. The highly hydrophilic
Gd-DTPA derivative is released into the aqueous phase. However, the created ONB lipid only
contains the weakly hydrophilic nitroso and benzaldehyde moieties in its headgroup, making
it a mostly hydrophobic molecule or a lipid with a small headgroup which will remain in the
lipid membrane. This can increase the liposome size in two ways. The ONB lipid with its
small headgroup prefers a smaller membrane curvature than the GDO lipid with the large
DTPA-Gd headgroup.[90-91] In addition, the ONB lipid facilitates liposome fusion because
its headgroup is smaller and less charged than that of the GDO lipid, both of which lower the
barrier for membrane fusion. This proposed mechanism explains the gradual increase in size
because an increasing number of GDO lipids is converted into ONB lipids as the light
irradiation proceeds (i.e., conversion increases with irradiation time and intensity), and the
more ONB lipids have been generated, the more they promote the increase in liposome size.
We next set out to determine whether irradiation of GDO liposomes is accompanied by the
20

release of encapsulated drug, allowing control of drug release from the vesicle by
photo-triggering.
c
a
b
d
Size (nm)
Size (nm)
e
Liposome photolysis
+
UV
Gd-DTPA-ONB
ONB lipid
Hydrophilic
head
GDO liposome
Prefers a smaller membrane curvature
Increase the liposome size
Increase the size
Lower the barrier for membrane fusion
Liposomes fusion
Figure 1. The effect of UV irradiation on GDO liposome size and morphology. (a) The
change of the particle size of a control sample of empty GDO liposome over 5 days in PBS
(without irradiation). (b) The change of the particle size of empty GDO liposome exposed to
21

UV light for increasing times (as indicated on the graph). (c) TEM micrograph of control
GDO liposomes (irradiation). (d) TEM micrograph of empty GDO liposomes exposed to UV
light (after 120 min of irradiation). (e) Proposed mechanism of GDO liposome enlargement
during photolysis.
3.4. Characterization of pH- and photo-triggered drug release from DOX-loaded GDO
liposomes (DOX/GDO liposomes)
3.4.1 Photo-triggered drug release from DOX/GDO liposomes
To investigate the extent and mechanism of release from drug-loaded GDO liposomes, we
used doxorubicin hydrochloride (DOX) as a model drug compound. We chose DOX because
it is a well-studied cancer drug with the added advantage of being fluorescent. Thus,
fluorescence imaging can be utilized as a secondary detection method for evaluating the MRI
imaging function of GDO liposomes in vivo. We first measured the encapsulation efficiency
at different ratios of GDO/DOX (see Figure S19 in the Supporting Information). At a ratio of
3:1, the encapsulation rate was 68.34%, higher than that for most reported drug loading
systems.[92-93]
Next, we determined the relationship between the size of the DOX-loaded liposomes and the
time of exposure to UV light (Figure 2a). With increasing exposure time (0–30 min), the
diameter of the liposomes progressively grew from the initial 260 nm (TEM data; Figure 2c
and Figure S20) to 700 nm, and the size distribution broadened. At longer exposure times, the
size of the nanoparticles rapidly decreased and eventually dropped to around 100 nm. At the
same time, a precipitate appeared in the UV-irradiated samples (Figure 2b) but not in control
samples without irradiation (even after 3 days; see Figure S21 in the Supporting Information).
The rapid appearance of red (DOX-containing) precipitate in the UV-irradiated samples
suggested that UV light, as expected, is a suitable trigger to accelerate drug release from
DOX/GDO liposomes. Visual inspection of the samples suggests that longer irradiation times
correspond to larger amounts of DOX being released, as judged by the increase in
DOX-containing sediment and the lighter color of the supernatant.
22

To quantify the relationship between illumination time and DOX release, DOX-loaded GDO
liposomes irradiated for various times were dialyzed and DOX release measured using DOX
fluorescence. Figure 2f plots the cumulative DOX release over time in 0.01 M PBS buffer
(pH=7.4) for the irradiated liposomes as well as a control sample of free DOX. Increasing
time of UV illumination gradually increased not only the cumulatively released DOX (e.g.,
from 45.2% to 66.4% after 48 h, when DOX release had plateaued for all samples), but also
the rate of release (i.e., the slope at 0–6 h in Figure 2f; see inset). In all samples, release from
the liposomes is slower than for free DOX. This slower or sustained release is actually a
desirable feature, because it allows time for the liposomes to reach the targeted tumor tissue
before they release DOX, reducing side effects.
Figure 2g plots the release efficiency for the irradiated liposome samples, i.e., their
cumulatively released DOX amount normalized to that for free DOX, at selected time points.
At the final time point (48 h), release efficiency is increased from 61.7% (no irradiation) to
92.2% (120 min irradiation) by photo-treatment. Interestingly, most of this impressive
increase in release efficiency occurs between 20 to 40 min of UV irradiation time (i.e.,
around the point where the change in size of the liposomes switches from growth to
shrinkage). Taking into account the observed size changes of DOX/GDO liposomes and the
proposed photocleavage mechanism of empty GDO liposomes, we hypothesize that two
competing processes occur during irradiation of DOX/GDO liposomes. One is the growth of
liposomes by lipid cleavage and vesicle fusion as observed for the empty liposomes. The
release data shows that some DOX release occurs at this stage, possibly from leakage during
fusion events. The second process is the formation of a DOX-containing precipitate which
likely also contains the ONB lipid (because no liposomal structures remain after 120 min of
irradiation). At short UV irradiation times (0–30 min), liposomes gradually grow and fuse
(accompanied by release of loaded DOX) as most of the produced ONB lipids are still
accommodated in the expanding liposomes. At longer times (UV exposure for >30 min),
much of the GDO lipid has been cleaved and larger quantities of DOX have been released. At
23

this point, the process of formation of the DOX-containing precipitate becomes dominant,
resulting in the reduction of the diameter of the liposomes because ONB lipid is sequestered
into the growing amount of precipitate.
Support for this hypothesis comes from the observation (by TEM) that the morphology of the
material suspended in the supernatant (Figure 2d) and found in the sediment (Figure 2e) is
similar in the sample exposed to UV light for 120 min. A comparison with a sample of
DOX/GDO liposomes before irradiation (Figure 2c) suggests that essentially all liposomes
have been dissociated after 120 min of UV exposure, forming a DOX-containing precipitate.
Sub-µm-size particles of the precipitate (a few 100 nm in size, see Figure 2d) remain
suspended in the supernatant. Taken together, the data show that adjusting the duration of
exposure to UV light can be used to control the release efficiency of DOX from DOX/GDO
liposomes. While we only investigated exposure time, other tunable parameters of the
photo-treatment (such as light intensity) will permit similar control over the drug
release.[94-95]
24

Figure 2. Photo- and pH-triggered drug release from DOX/GDO liposomes in vitro. (a) The
change of the particle size of DOX/GDO liposomes with time of exposure to UV light. The
size of the particles in solution initially increases, then decreases to below the size before
irradiation. (b) Macroscopic appearance of DOX/GDO liposome samples after different times
of exposure to UV light. A DOX-containing (red) precipitate forms, with the amount of
precipitate increasing with irradiation time. (c) TEM micrograph of a DOX/GDO liposome
prior to UV irradiation. (d) TEM micrograph of the supernatant of a suspension of
DOX/GDO liposomes irradiated for 120 min. (e) TEM micrograph of the sediment from a
suspension of DOX/GDO liposomes irradiated for 120 min. (f) Cumulative DOX release
25

from DOX/GDO liposome suspensions irradiated for varying times (0–120 min) with UV
light, plotted against dialysis time. (g) DOX release efficiency of DOX/GDO liposome
suspensions irradiated for varying times (0–120 min) with UV light at various dialysis times,
calculated by normalizing the data from part (f) to the amount of DOX released from the
sample of free DOX at 48 h of dialysis. (h) Cumulative release of DOX for free DOX,
DOX/DOPC, and DOX/GDO liposomes irradiated for different times (0, 5, and 20 min) at
four pH values: pH=8.5 (control), pH=7.4 (physiological pH), pH=6.5 (tumor extracellular
environment), and pH=4.2 (late endosome/lysosome). (i) The DOX release efficiency
(cumulative release normalized to that of free DOX at the same pH) for the same samples as
in part (h) at 48 h. For parts f and h: (*: p＜0.05; ***: p＜0.001).
3.4.2 pH-triggered drug release from DOX/GDO liposomes
We investigated drug release from DOX/GDO liposomes in phosphate buffers at several pH
values relevant to different cellular and in vivo environments: a slightly alkaline pH (pH=8.5;
as the control), physiological pH (pH=7.4), the slightly acidic pH of the tumor extracellular
environment (pH=6.5), and the low pH of late endosomes/lysosomes (pH=4.2). Free DOX
and DOX-loaded liposomes prepared from the neutral, pH-stable phospholipid DOPC
(DOX/DOPC) served as the controls. At each pH, the DOX/GDO liposomes were irradiated
with UV light for 0 (no irradiation), 5, or 20 min to assess their capability for photo-triggered
release at the different pH values. The cumulative release of DOX over time is plotted in
Figure 2h. The amount of DOX released from DOX/DOPC liposomes is between that
released from DOX/GDO liposomes without and with short UV irradiation at the highest pH
(pH=8.5). At the other, lower pH values, the amount of DOX released from DOX/DOPC
liposomes is lower than that released by any GDO liposomes (with or without UV irradiation)
with the difference in DOX release increasing as pH decreases. This is consistent with the
interpretation that the GDO lipids are cleaved under acidic conditions, leading to DOX
release. For DOX/GDO liposomes, the overall trend is that DOX release increased with
decreasing pH as well as with exposure time to UV light. However, the cumulative release
from samples containing only free DOX also increases strongly with decreasing pH because
26

of changes in DOX hydrophobicity and thus solubility with pH. Thus, Figure 2i shows the
release efficiency at the different pH values, i.e., the amount of DOX (after 48 hours) released
from liposomes normalized to the amount released from the free DOX sample. While the
cumulative release of DOX (and its absolute change with UV irradiation) was lowest in an
alkaline environment (pH = 8.5, used as a control), photo-release efficiency was highest at
this pH (increased from 57.1% without irradiation to 90.3% with 20 min irradiation). The
slight positive slope in the cumulative release at this pH may be due to alkaline hydrolysis of
the ester bond (see also Figure S12 in the Supporting Information). For the samples at
physiological (pH=7.4) and acidic pH, the release efficiency with and without irradiation
increases with decreasing pH, likely due to an increase in the acid-catalyzed cleavage of the
GDO lipid. The relative increase in release efficiency with UV irradiation (e.g., for 20 min) is
very similar for these three samples, increasing only slightly from 10.1% (at pH=7.4) to 11.7%
(pH = 6.5) and 13.8% (pH=4.2).
The increased DOX release at low pH compared to pH=8.5 and pH=7.4 demonstrates that
GDO liposomes can intelligently distinguish the tumor environment from normal cells due to
their pH-controlled release ability. Light-triggered release then provides a further tool for
precise regulation of drug release at the site of the lesion. The increased release at the pH of
late endosomes and lysosomes has great potential for future development of the GDO
liposomes, where addition of a peptide ligand could add receptor-based targeting and
increased endosomal uptake into target cells.[96-99]
3.5 Cytotoxicity of empty and DOX-loaded GDO liposomes
We assessed the cytotoxicity of empty GDO liposomes in macrophages (Raw 264.7) and
three cancer cell lines (MCF-7, Hela7702 and HepG2) to determine an appropriate
concentration for experiments with drug-loaded liposomes. As shown in Figure 3a, the GDO
liposomes were essentially nontoxic to macrophages throughout the investigated
concentration range, but showed high toxicity toward cancer cells at higher concentrations,
with an onset of toxicity (in HepG2 cells) at 1.25 mM. We also irradiated the four cell lines
27

with UV light (365 nm at 60 W) for 0–20 min (after 24 h of culture) to determine an upper
time limit for UV irradiation. As shown in Figure 3b, exposing cells to UV light for 5 or 10
min did not result in significant cell death. However, after 20 min of irradiation, viability of
all four cell lines was reduced by 15–20%. Based on the results of these preliminary
experiments, we added DOX-loaded liposomes at a concentration of 1 mM of GDO lipid to
the four cell lines and irradiated them with UV light for 5, 10, or 20 min to assess their
light-triggered release capability. After 24 h, the viability of cells treated with the
DOX-loaded liposomes without UV irradiation was about 65%, and viability decreased
further with increasing time of UV irradiation. After 20 min of irradiation, cell viability
dropped to the level observed for cells treated with an equivalent concentration of free DOX
(about 38%). This indicates that the released amount of DOX can be controlled by UV
irradiation, and that more DOX is released at longer irradiation times. Figure S35, showing
confocal microscopy images of free DOX, DOX/GDO and DOX/GDO-UV added to MCF-7
cells provides a visual comparison of their effects.
While these in vitro experiments show that free DOX enters cells quickly and is slightly more
cytotoxic than liposomal DOX, the therapeutic efficacy of free DOX is highly limited by side
effects caused by its lack of targeting and indiscriminate toxicity such as cardiotoxicity and
myelosuppression.[100] To determine whether loading DOX into GDO liposomes enables a
more targeted approach that will reduce toxicity and improve drug efficacy, we turned to the
MRI capabilities of the GDO liposomes.
28

Figure 3. Cytotoxicity of empty and DOX-loaded GDO liposomes. (a) Cytotoxicity (MTT
assay) of empty GDO liposomes versus concentration, against three cancer cell lines (MCF-7,
Hela7702 and HepG2) and macrophages (Raw 264.7). (b) The effect of UV irradiation, free
DOX, and 1 mM DOX-loaded GDO liposomes with and without UV irradiation on the
viability of three cancer cell lines (MCF-7, Hela7702 and HepG2) and macrophages (Raw
264.7).
3.6. T1-Magnetic resonance imaging of DOX/GDO liposomes in vitro and in vivo
The longitudinal relaxation time (T1) and transverse relaxation time (T2) of protons are the
main parameters that affect the quality of MRI.[101] Gd complexes are a typical T1-MRI
agent which mainly improves the longitudinal relaxation efficiency r1 by affecting the
longitudinal relaxation time T1 of the water protons to obtain a brighter and higher-contrast
MR image. Generally, the relaxivity of an MRI contrast agent is determined by the following
formula:
1
1
ꢐ ꢑ
+ ꢏ CA
ꢅ
=
ꢉ
ꢅ
ꢉ
ꢅ,ꢊꢋꢌꢌꢍꢎ
Here, T1 is the longitudinal relaxation time in presence of the contrast agent, and T1,tissue is the
ꢐ
ꢑ
longitudinal relaxation time of the solvent or tissue without contrast agent. CA is the
concentration and r1 is the longitudinal relaxivity of the contrast agent. A similar equation
can also be applied to the transverse relaxivity r2.
29