Trigger Responsive Dendrimer- Liposome Hybrid Nanocarrier System
for Targeted Drug Delivery
Frehiwot Gebrehiwot
RET 2012
NSF-RET Grant EEC 0743068 and RET EEC 1132694
Laboratory for Polymeric Nanomaterials in Biological Sciences
Department of Biopharmaceutical Sciences
Mentors/Advisors: - Dr. Seungpyo Hong, Principal Investigator
- Suhair Sunoqrot, Ph.D. candidate
- Eri Iwasaki, M.S. candidate
Introduction
Cancer remains a devastating disease with nearly 10 million new cases worldwide every year1. It is
projected that in the United States alone 1,638,910 people will be diagnosed with cancer in 2012 and
577,190 of them will die1 from it. Chemotherapeutic drugs that are currently used in cancer treatments
often also kill healthy cells and cause severe side effects. Therefore it is desirable to develop therapeutic
methods that can either passively or actively target cancerous cells. Passive targeting relies on the
characteristic features of tumors which are the leaky blood vessels and poor lymphatic drainage. This is
known as the enhanced permeability and retention (EPR) effect2, which promotes accumulation of
nanocarriers within 50-200 nm in tumor sites3. However, passive targeting suffers from several
limitations because some drugs cannot diffuse efficiently and is an uncontrolled process. On the other
hand, active targeting employs conjugating nanocarriers containing chemotherapeutic drugs with
molecules that bind to overexpressed receptors on the target cells surface4. However, surface exposure
of targeting ligands can also lead to increased unwanted uptake and rapid clearance5. This study will
focus on achieving a high targeting efficacy by taking advantage of both passive and active targeting
methods.
Figure 1: Schematic representation of different mechanisms of drug delivery by nanocarriers to
tumours. Passive targeting is achieved through the increased permeability of the tumour vasculature and
ineffective lymphatic drainage (EPR effect). Active targeting is achieved through cell-specific
recognition and binding by ligands4.
Polyamidoamine (PAMAM) dendrimers have shown promise for biomedical applications due to their
well-defined structure, high water solubility and multifunctionality. On the other hand, liposomes have
several advantageous characteristics including biocompatibility, high loading capacity for hydrophilic
molecules and physical encapsulation of cargos without the conjugation chemistries6. The goal of this
study is to design a trigger responsive dendrimer-liposome habrid nanocarrier, or nanohybrid, system.
Folate (FA) targeted dendrimers will be encapsulated within liposomes to create the hybrid system. In
order to take advantage of both passive and active targeting, the dendrimers must be released from the
liposomes before internalization. However, internalization of liposomes is likely to occur through
endocytosis or other nonspecific pathways7. Therefore, we hypothesize that dendrimer release from
liposome can be achieved by using pH-sensitive liposomes programmed to release their payload at the
pH of the tumor microenvironment (pH 6.7 compared to the physiologic pH 7.4)8.
Materials and Methods
Materials
Generation 4 (G4) PAMAM dendrimer, N-hydroxysuccinimide-rhodamine B (NHS-RHO), folic acid
(FA), cholesterol, methanol, and dichloromethane (DCM) were all obtained from Sigma-Aldrich (St.
Louis, MO). 1,2-dihenarachidoyl-sn-glycero-3-phosphocholine (21PC), 1,2-distearoyl-sn-glycero-3phosphoethanolamine-N-(amino(polyethylene glycol)-2000) (DSPE-PEG 2000), and 1,2-dimyristoylsn-glycero-3-phosphocholine (DMPC) were purchased from Avanti Polar Lipids Inc. (Alabaster, AL).
Preparation of G4 PAMAM Dendrimer Conjugates
G4 PAMAM dendrimers were fluorescently labeled by conjugation with NHS-RHO as previously
described9. Briefly, amine-terminated G4 was dissolved in 4 mL of sodium bicarbonate buffer (pH 9.0),
to which 500 µL of NHS-RHO in DSMO was added, and the reaction mixture was vigorously stirred at
room temperature (RT) for 24 hours. Unreacted NHS-RHO was removed by membrane dialysis using a
Spectra/Por dialysis membrane (MWCO 3500, Spectra Laboratories Inc., Rancho Dominguez, CA) in
excess deionized distilled water (ddH2O) for 2 days. The purified G4-RHO-NH2 conjugates were
lyophilized over 2 days using a Labconco FreeZone 4.5 system (Kansas City, MO) and stored at -20 0C.
Next, FA was conjugated to G4-RHO-OH. FA was activated by EDC and NHS in 1.5 mL of DMSO
through vigorous stirring at RT for 1 hour. The activated FA solution was added to 20 mg of G4-RHOOH in 1 mL of ddH2O, followed by membrane dialysis as described above, resulting in G4-RHO-FAOH.
Encapsulation of Dendrimers into Liposomes
Unilamellar liposomes were prepared using a film hydration method followed by extrusion as previously
described10. pH-sensitive liposomes were composed of 21PC (5.7 mg, 6.5 x 10-6 mol), DSPA (1.6 mg,
2.2 x 10-6 mol), cholesterol (0.3 mg, 0.6 x 10-6 mol), and DSPE-PEG2000 (1.4 mg, 0.5 x 10-6 mol).
Regular liposomes were composed of DMPC (4.5 mg, 6.7 x 10-6 mol), cholesterol (1.1 mg, 2.9 x 10-6
mol), and DSPE-PEG2000 (1.4 mg, 0.5 x 10-6 mol). First, lipids (10 µmol total) were dissolved in 5 mL
of DCM in a round-bottom flask. The flask was connected to a rotary evaporator (Rotavapor RII, Buchi,
Switzerland) at 50 0C for 1 hour to evaporate DCM until completely dried.
The dried lipid films were hydrated in 1 mL of either 0.1 mg/ml FA-targeted dendrimer or 0.1 mg/ml
non-targeted dendrimer solution in ddH2O, followed by vortexing for 15 minutes to form multilamellar
liposomes. Multilamellar liposomes were sonicated in a bath sonicatior for 30 minutes and then extruded
20 times through a polycarbonate membrane of 100 nm pore size using a Lipofast Pneumatic extruder
(Avestin Inc., Ottawa, Canada). The resulting unilamellar liposome suspension was centrifuged at
20,000 rpm for 1 hour to remove residual dendrimers. The pellet was resuspended in 1 mL of 5%
sucrose, lyophilized over 2 days and stored at -20 0C for further characterization.
Results
Characterization of FA-targeted dendrimer-liposomes nanohybrids
Particle size (diameter in nm) and surface charge (zeta potential in mV) of liposomes were obtained
from quasi-erastic laser light scattering using a Nicomp 380 Zeta Potential/Particle Sizer. (Particle
Sizing Systems, Santa Barbara, CA). The measurements were performed using samples suspended in
ddH2O at concentration of 100 µg/mL.
Figure 2: Particle size of pH-sensitive liposomes.
Figure 3: Particle size of regular liposomes.
The encapsulation process yielded pH sensitive liposome-dendrimer hybrids with a controlled size of
126 nm in diameter and a charge of -22.21 mV (Figure 2). The size and zeta potential of regular
liposome-dendrimer hybrids was also measured with particle size and zeta potential of 221 nm and 28.73 mV respectively (Figure 3). Particle size measured over 700 nm is attributed to aggregation of
liposomes
Loading Efficiency
Loading was defined as the dendrimer conjugate content in the liposomes. Loading was determined by
dissolving 3.1 mg of lyophilized liposomes in 500 µL of 0.1% Triton x-100, followed by filtration
through a 0.45 µm syringe filter. The fluorescence intensity from the filtrates was then measured using a
SpectraMAX GeminiXS microplate spectrofluorometer (Molecular Devices, Sunnyvale, CA). The
amount of G4-RHO-FA-OH released was determined from the standard curve of G4-RHO-FA-OH
fluorescence versus concentration in 0.1% Triton X-100. Loading efficiency was defined as the ratio of
the actual loading obtained to the theoretical loading.
Table 1. Loading efficiency of
G4-RHO-FA-OH-liposomes nanohybrids
Fluorescence (A.U.)
Loading Efficiency (%)
pH-sensitive
Liposomes
Regular
14.659
16.690
5.4
Liposomes
6.3
Figure 4: Standard curve of G4-RHO-FA-OH
fluorescence versus concentration in 0.1%
Triton X-100.
The calculated loading efficiency for pH-sensitive liposomes and regular liposomes were 5.4% and
6.3% respectively (Table 1). Both loading values were much lower than expected (~80%). We believed
that the charge of the dendrimers which was neutral was affecting the encapsulation process.
In order to determine that the low loading of dendrimers into liposomes was due to the charge of
dendrimers, we encapsulated positively charged dendrimers (G4-RHO-NH2 ). The same protocol was
used to prepare the liposomes. Following, the particle size and surface charge of the liposomes was
obtained (Figure 4). The size and zeta potential of G4-RHO-NH2-liposome nanohybrids were 176.8 nm
and -31.27 mV, respectively. The obtained size was bigger than the size of G4-RHO-FA-OH-liposomes
nanohybrids (124.7 nm). This could be explained by the positively charged dendrimers inducing
aggregation with negatively charged vesicles leading to higher mean diameter.
Figure 5: Particle size of G4-RHO-NH2liposome nanohybrids.
The loading efficiency of the nanohybrids with the positively charged dendrimers was obtained
following the previously described method. Fluorescence of the pellet was measured to be 142.55 A.U.
and the loading efficiency was calculated to be 28.3%. Encapsulation efficiency of G4-RHO-FA-NH2
was 4.5 times higher than that of G4-RHO-FA-OH (6.3%). As these results show it seems that
encapsulation efficiency depends on the charge of dendrimers.
Future Plan
It is evident that the encapsulation efficiency depends on the charge of dendrimers. The next stage in the
study is to investigate different components and optimal ratio of lipids to maximize the encapsulation
efficiency.
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