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Experimental Therapeutics |
Department of Biomedical Engineering, Duke University, Durham, North Carolina 27708 [D. E. M., G. A. K., A. C.], and Departments of Radiation Oncology [M. W. D.] and Radiology [M. R. Z.], Duke University Medical Center, Durham, North Carolina 27710
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ABSTRACT |
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 Top ABSTRACT INTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONREFERENCES |
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2-fold increase in tumor localization
compared to the same polypeptide without hyperthermia. We observed
aggregates of the thermally responsive ELP by fluorescence
videomicroscopy within the heated tumor microvasculature but not in
control experiments, which demonstrates that the phase transition of
the thermally responsive ELP carrier can be engineered to occur
in vivo at a specified temperature. By exploiting the
phase transition-induced aggregation of these polypeptides, this method
provides a new way to thermally target polymer-drug conjugates to solid
tumors. |
INTRODUCTION |
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TopABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONREFERENCES |
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Soluble macromolecular carriers are useful drug delivery vehicles because they increase the plasma half-life of low molecular weight drugs, increase the solubility of hydrophobic drugs, provide passive targeting to tumors by the enhanced permeability and retention effect (6) , and can also permit controlled release of the drug through the use of degradable linkers (5) . Because vascular permeability increases at temperatures between 40°C and 45°C, hyperthermia treatments can also enhance the delivery of drugs to solid tumors (7) . Furthermore, when combined with chemo- and radiotherapy, hyperthermia can synergistically enhance tumor cytotoxicity (8 , 9) .
Here, we propose a novel thermal targeting scheme using a thermally
responsive, genetically engineered
ELP.3
ELPs are biopolymers of the pentapeptide repeat Val-Pro-Gly-Xaa-Gly,
where the "guest residue" Xaa can be any of the natural amino acids
except Pro (10)
. ELPs undergo an inverse temperature phase
transition; they are soluble in aqueous solutions below their
Tt, but they hydrophobically collapse and
aggregate at temperatures greater than Tt
(10
, 11)
. We hypothesized that an ELP with a
Tt intermediate between Tb
and Th would enable thermally targeted drug
delivery to a locally heated region. In this scenario, the ELP would be
soluble systemically because its inverse Tt
(Tt 
41°C) is greater than
Tb (Tb 37°C–38°C),
but it would become insoluble and accumulate in locally heated regions
where the temperature was increased above Tt by
externally targeted hyperthermia (Th
42°C–43°C). In concept, this method synergistically combines
thermal targeting through the ELP phase transition with the established
advantages of both polymeric carriers (e.g., increased
plasma half-life, high loading capacity) and hyperthermia as an
anticancer treatment modality (e.g., increased sensitivity
to therapeutics, greater macromolecular tumor extravasation).
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MATERIALS AND METHODS |
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TopABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONREFERENCES |
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ELP Synthesis.
Our approach to the synthesis of ELPs has been described elsewhere
(13)
. Briefly, synthetic genes encoding the two different
ELP sequences were constructed as follows (Fig. 1)
. Short gene segments (10 pentapeptides for ELP1 and 16 pentapeptides
for ELP2) were assembled by annealing chemically synthesized
oligonucleotides (Integrated DNA Technologies, Coralville, IA) encoding
the sense and antisense strands to form a gene cassette, which was then
ligated into pUC19 (New England Biolabs, Beverly, MA). These DNA
segments were oligomerized by a process we term "recursive
directional ligation," a convenient and flexible method to rapidly
assemble a specified number of tandem gene repeats in a defined
orientation.4
The final gene encoded 150 ELP
pentapeptides for ELP1 and 160 pentapeptides for ELP2. The oligomerized
genes were then excised from pUC19 and ligated into a modified pET25b
expression vector (Novagen, Madison, WI). The expression vector
contained translation initiation and termination codons and the codons
for short leading (Ser-Lys-Gly-Pro-Gly) and trailing (Trp-Pro)
sequences. Standard molecular biology protocols were used for all DNA
manipulations (14)
.
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Characterization of the Inverse Transition.
The inverse temperature phase transition of the ELPs was characterized
by monitoring the absorbance at 350 nm as a function of temperature on
a UV-visible spectrophotometer equipped with a multicell thermoelectric
temperature controller (Cary 300 Bio; Varian Instruments).
Reversibility of the transition was examined by first heating an ELP
solution in PBS, typically from 20°C to 60°C at a rate of
1°C/min, and then cooling the solution to 20°C at the same rate.
The Tt was defined from the heating profile as
the temperature at 5% of maximum turbidity.
Labeling.
The ELP constructs have two primary amines (the
NH2 terminus and a single lysine at the second
residue). Labels for the in vivo experiments were conjugated
to these amines using succinimidyl ester chemistry. Rhodamine Red-X
succinimidyl ester (Molecular Probes, Inc.) was used for the
fluorescence videomicroscopy experiments, and SIB (17)
was
used for the radiolabeling experiments, where the iodine was either
131I, 125I, or
127I (nonradioactive). For paired label
biodistribution studies, ELP1 was conjugated to
[131I]SIB, and ELP2 was conjugated to
[125I]SIB.
In a typical conjugation reaction, ELP1 or ELP2 was dissolved in 100
mM sodium bicarbonate (pH 8.4). The reporter, which was
dissolved at a concentration of 10 mg/ml in dimethylformamide, was
slowly added while mixing to a final molar excess of 4 for
rhodamine-to-ELP coupling and 40 for SIB-to-ELP coupling. The
reactants were incubated with continuous stirring for 2 h at room
temperature. Insoluble matter was removed by centrifugation at 4°C.
The ELP was then purified from soluble free label by inverse transition
cycling. Labeling yield was calculated using UV-visible
spectrophotometry (UV-1601; Shimadzu Scientific Instruments, Inc.).
Typically, the molar label:ELP ratio was 0.5 for rhodamine and
1.0 for SIB.
Serum Stability.
The rhodamine-ELP and SIB-ELP conjugates were separately diluted 1:6 in
fresh, heparinized mouse serum. This dilution ratio approximates that
of the in vivo experiments. The ELP-serum mixture was then
incubated at 37°C for up to 2 days, and aliquots were removed for
analysis at various times points for analysis by SDS-PAGE, which
separated the conjugated and free label into distinct bands. The
conjugates were visualized under UV transillumination and quantitated
by scanning densitometry. Free and conjugated SIB were separated by
instant TLC, and each fraction was counted for radioactivity.
In Vivo Fluorescence Videomicroscopy.
Athymic mice were implanted with dorsal skin fold window chambers
containing a small piece of tumor tissue (0.1
mm3
human ovarian carcinoma; SKOV-3) placed near
the center of the window (18, 19, 20)
. The preparations were
used 7–8 days after implantation, when the tumors had grown to 2–3 mm
in diameter. The mice were anesthetized with sodium pentobarbital (80
mg/kg, i.p.) and positioned laterally recumbent on a microscope stage
(Carl Zeiss, Inc., Thornwood, NY). The window chamber, which was
connected to a temperature-controlled water bath that had been
previously calibrated to the temperature within the chamber (7
, 21)
, was maintained either at the physiological s.c. temperature
in mice, 34°C (7)
, or at 42°C. A region of the tumor
containing clearly visible tumor microvasculature was selected under
transillumination with a x20 objective. Then, under epi-illumination
using a dichroic filter set for rhodamine, images were acquired by a
SIT camera (Hamamatsu C2400-08) and recorded in S-VHS format
(Mitsubishi BV-1000). In a typical experiment, 200 µl of 100
µM rhodamine-labeled ELP in PBS (1.2 mg/mouse) were
injected into the cannulated tail vein. Three groups of five animals
each were studied: (a) ELP1 at 42°C; (b) ELP2
at 42°C; and (c) ELP1 at 34°C. Images were recorded
continuously for 40 s after injection and for 10 s every 2
min for 50 min.
Image Analysis of in Vivo Fluorescence
Videomicroscopy.
Accumulation of the rhodamine-labeled ELPs in the tumors was quantified
by digital analysis of the videomicrographs as follows. A TIFF image
was created for each of 33 time points (10 s preinjection; 30 s, 1
min, and 2 min after injection; and every 2 min thereafter up to 50
min) by averaging 30 frames over 3 s of footage using Scion Image
v. 1.62 (NIH Image, as modified by Scion Corp.). The images were
analyzed for total window fluorescence intensity and for fluorescence
intensity subdivided in the vascular and interstitial spaces within the
total window. For total window intensity, the average pixel value was
calculated for the entire window (440 x 515 µm). The
background (the 10 s preinjection time point) was subtracted from
each of the other time points, which were then normalized to the
initial time point (30 s after injection). For calculating the relative
vascular and interstitial intensities, regions in the 30 s
postinjection videomicrograph were defined as either vascular or
interstitial. These definitions were determined by numerically
computing the gradient of the intensity at each point (i.e.,
pixel) in the window and then using a threshold parameter to define the
edges of vessels. The average intensity of these edges was calculated
for subregions of the image. Pixels within the subregion that had an
intensity greater than the local average for the edges and the edges
themselves were classified as vascular. These spatial definitions were
then used as a mask for later time points, allowing pixel values to be
averaged independently for the interstitial and vascular regions. The
average intensities for the vascular and interstitial areas were
analyzed by subtracting the background and normalizing the initial
vascular average intensity.
Tumor Localization Studies.
Athymic mice (Balb/C nu/nu) were inoculated s.c. on the
lateral thigh with 50 µl of tumor homogenate (human glioma; D-54MG).
The experiments were performed 7–10 days later, when the tumors had
grown to at least 3 mm in diameter. The mice were anesthetized with
sodium pentobarbital (80 mg/kg, i.p.), and the tumor-bearing leg was
taped into a wire jig that allowed each mouse to be suspended such that
only the leg was immersed in a temperature-controlled water bath
maintained at 43.7°C (22)
. There were two groups of five
animals each; one group received the hyperthermia treatment, whereas
the other, normothermic control group was left at room temperature. At
the beginning of the experiment, [125I]SIB-ELP1
and [131I]SIB-ELP2 (3 µCi of each) were
coinjected into the tail vein of each mouse. The radiolabeled ELPs were
supplemented with ELPs conjugated to nonradioactive SIB so that the
total dose injected was 300 µg of each ELP per animal (each ELP at
50 µM in 100 µl of PBS). After 45 min, both
groups were euthanized with an overdose of halothane (Halocarbon
Laboratories, River Edge, NJ). The tumors were resected and counted for
radioactivity on a LKB 1282 dual-channel gamma counter (Wallac), with
crossover correction for 131I spillover into the
125I counting window.
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RESULTS AND DISCUSSION |
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41°C. This Tt
was chosen because it is greater than the normal body temperature but
lower than temperatures that can be readily achieved clinically by
focused hyperthermia methods (23)
. We hypothesized that
ELP1 would be systemically soluble below Tt, but
would become insoluble and aggregate in the target region where the
temperature was greater than Tt. We controlled
the Tt by adjusting the hydrophobicity of the
amino acid sequence (12)
. For ELP1, the
Tt was increased to 41°C from 27°C, which is
the Tt of the natural elastin repeat VPGVG
(12)
, by substituting Ala and Gly for the more hydrophobic
Val at the fourth position of the pentapeptide. The thermally
unresponsive control, ELP2, was designed to have a
Tt greater than Th by
incorporating a greater fraction of hydrophilic Ala and Gly residues so
that it would remain soluble even in the heated tissues.
In Vitro Characterization of ELPs.
The inverse transition behaviors of both ELP1 and ELP2 were
characterized by monitoring solution turbidity as a function of
temperature. Fig. 2
shows a heating and cooling turbidity profile for the ELP1 carrier.
Below the Tt, the polypeptide solution was clear,
but upon further heating, the solution became turbid because of ELP
aggregation. The rapid increase in turbidity upon reaching
Tt shows that the transition is very sharp with
respect to temperature, occurring over a range of less than 2°C. The
inverse transition of each ELP was completely reversible, and no end
point hysteresis was observed. The slight difference between the paths
of the heating and cooling traces is due to the slower kinetics of
disaggregation as compared with aggregation. From the heating turbidity
profiles, the Tt is defined as the temperature at
the onset of turbidity (5% maximal turbidity) to precisely
characterize the first appearance of ELP aggregates.
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also shows that conjugation of SIB and rhodamine reduced the
ELP1 Tt by 4°C. We hypothesize that this
shift is caused by hydrophobic interactions facilitated by the close
proximity of the conjugated, nonpolar labels and the ELP chain,
analogous to the decrease in the Tt observed with
increasing guest residue hydrophobicity. Loss of the hydrophilic,
charged amine groups on conjugation may similarly contribute to the
Tt reduction. The decrease in
Tt after conjugation has important implications
for the loading of ELPs as therapeutic carriers, particularly for drugs
with a significant hydrophobic character. The decrease of
Tt by 4°C observed here does not require any
compensation other than an adjustment of the ELP plasma concentration
during thermally targeted delivery (discussed below). However, for
drugs that are more hydrophobic or for higher drug:carrier loading
ratios, the carrier can be designed to have a higher
Tt in anticipation of a larger downward shift
caused by conjugation of the drug. We also studied the effect of mixing ELP1 and ELP2 on their inverse transition behaviors. This has direct relevance to the tumor biodistribution studies carried out with ELP1 and ELP2, which were conjugated to [131I]SIB and [125I]SIB, respectively, mixed in equimolar proportions, and injected. Coinjected paired labels are desirable because they enable the simultaneous study of the tumor localization of both the thermally responsive (ELP1) and thermally unresponsive (ELP2) carriers in the same animal, thereby minimizing the effect of animal-to-animal physiological variability. To use paired labels, however, it was critical to understand whether ELP1 and ELP2 in mixture would each exhibit independent aggregation at their respective Tts or whether they would display a cooperative transition behavior that would preclude paired label biodistribution studies using coinjected ELP1 and ELP2.
The inverse transition behavior of an equimolar ELP1 and ELP2 was
monitored by a solution turbidity assay in PBS supplemented with 1
M NaCl (Fig. 3)
. (1 M NaCl was added to depress the
Tt so that the transitions of both components in
the mixture could be observed in an experimentally convenient
temperature range.) The mixture showed biphasic aggregation behavior:
on increasing the temperature, two distinct aggregation events were
observed that closely overlaid the aggregation profiles for ELP1 and
ELP2 obtained separately. We conclude that the
Tts for ELP1 and ELP2 are sufficiently different
such that interactions between the more hydrophilic ELP2 and the
relatively hydrophobic ELP1 are minimal, even after ELP1 has undergone
its transition to the collapsed state. These results clearly show that
ELP1 and ELP2 maintain independent aggregation behavior in solution,
indicating that paired label studies can be used with these carriers.
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. PBS with 0.9 mM BSA and murine plasma both
reduced the ELP1 Tt by 4°C compared with
distilled water.
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shows that the Tt is an inverse logarithmic
function of ELP concentration. Therefore, although serum cosolutes and
the conjugation of hydrophobic reporters and drugs act to
decrease the Tt for a given
concentration, the increase in Tt with
decreasing concentration can be exploited to compensate for these
effects. Achieving and maintaining the intended ELP plasma
concentration are important considerations for thermal targeting
because if the in vivo concentration is too high, the
Tt will drop below Tb, and
the carrier will undergo its phase transition systemically. This would
eliminate selective targeting to the heated region and could
potentially deliver an unintended dose of aggregated carrier to other
organs. Conversely, if the in vivo concentration is too low,
then the Tt would be greater than
Th, and the carrier will remain soluble
systemically. Although this would eliminate thermal targeting to the
heated tumor, this latter outcome is less problematic because there
would also be no thermally induced aggregation in nontargeted organs.
During therapy, Tt should therefore remain above
Tb but can safely be allowed to increase to above
Th (e.g., as the carrier concentration
decreases with time due to systemic clearance).
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10-fold range in ELP concentration (Fig. 5)
. If higher or
lower ELP doses are required for therapy, ELP sequences with different
compositions can be readily synthesized such that they exhibit a
Tt between 37°C and 42°C in a different range
of concentrations as compared with ELP1. We also note that if the total
dose of drug needs to be limited below a specified threshold as defined
by its systemic toxicity, free ELP carrier can be mixed with the
ELP-drug conjugate to provide the necessary polypeptide concentration
required to achieve the target Tt. Based on these results and experimental considerations, we selected a target plasma concentration of 5 µM for the radiolabel in vivo experiments. Turbidity profiles of the SIB-labeled ELP1 inverse transition in the PBS + 0.9 mM BSA solution indicated that this ELP concentration will yield in a Tt of 41°C in blood (data not shown). We then calculated the dose required to achieve this concentration using an estimated mouse plasma volume of 1 ml (25, 26) .
Serum Stability.
Before the in vivo experiments, we tested the stability of
the fluorescent (rhodamine) and radioactive (SIB) conjugates in serum
as a function of time. No ELP degradation was observed; however, both
labels were cleaved over time. No loss of SIB was observed in ELP
incubated in physiological saline (Fig. 6A)
, which suggests that the cleavage was due to an enzyme
present in the serum. Fig. 6A
also shows that the SIB-ELP
conjugate was stable over time, and 90% remained conjugated after
48 h in serum. The rhodamine label was cleaved from the ELP more
rapidly, with only 60% remaining conjugated after 24 h (Fig. 6B)
. However, over the 1-h time frame of the in
vivo experiments, loss of either label from the ELP carrier was
negligible (<2%).
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Fig. 7
shows representative videomicrographs acquired at 30 s and at 30
min after injection for all three groups of animals. Within 1–2 min
after the thermally responsive ELP1 carrier was injected into mice with
window chambers heated to 42°C, fluorescent particles were observed
(Fig. 7
, ii.). None were observed for ELP1 without
hyperthermia (Fig. 7
, iv.) or for ELP2 with hyperthermia
(Fig. 7
, vi.) at any time, strongly suggesting that these
particles are ELP aggregates resulting from the inverse temperature
transition. The aggregates often attached to the vessel wall, and they
grew in intensity and size over time. Once attached to the vessel wall,
the aggregates were typically stable throughout the experiment. More
rarely, as the particles grew larger, they were sheared from the vessel
walls and carried away by the blood flow. Because the inverse
transition is reversible, any particles washed from the heated tissue
would be expected to rapidly disaggregate and resolubilize. The
observation of ELP1 aggregates only in the heated window chambers is an
exciting finding because, to our knowledge, it is the first in
vivo demonstration that a thermal phase transition of a free
polymer in solution can be engineered to occur at a specified
temperature within a complex physiological system.
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. Over the first 30 min, the accumulation of ELP1 in the heated tumor
increased rapidly. For the remainder of the experiment, however, tumor
ELP levels remained constant. The reason for this apparent saturation
could be systemic clearance or local depletion of the ELP from the
circulation. Decreasing plasma ELP concentration could have caused the
Tt to increase above Th, as
described above, thereby halting any further accumulation. After 50
min, the total window intensity had increased 250% relative to the
initial window intensity at 30 s. This represents a nearly 70%
greater tumor accumulation of ELP1 in the heated tumor as compared with
ELP1 at 34°C and a 54% increase versus ELP2 at 42°C.
For all time points after 4 min, ELP1 delivery to the heated tumor was
significantly greater than that for both the unheated ELP1
(P < 0.05) and heated ELP2
(P < 0.1) control groups.
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Fig. 9
shows the interstitial and vascular fluorescence intensities
quantitated using this algorithm for ELP1 in the heated tumor. At the
initial 30 s time point, the interstitial intensity was measured
as 53% of the vasculature. Both regions increased in intensity over
the course of the experiment, although the interstitium increased more
rapidly, and, at 50 min, its intensity had increased to 86% of the
vasculature. Accumulation in the interstitial space is likely due to
extravasation of both soluble ELP and small aggregates, presumably via
endothelial cell gaps, and is perhaps enhanced by further aggregation
once extravasated. Thus, these results suggest that in addition to the
larger aggregates that accumulate in the vessels, extravasation of the
ELP also plays an important role in the total ELP accumulation observed
within the heated tumor. Studies on the mechanism via which the ELPs
interact with endothelial and tumor cells as a function of temperature
are currently in progress to elucidate the origin of this effect.
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shows that tumor localization of the thermally responsive ELP1 was 80%
greater for the animals that received the hyperthermia treatment than
for the same ELP in unheated tumors (P < 0.01). Unlike the window chamber experiment, the data for the high
temperature control were obtained from the same animals using
coinjected, paired radiolabels. For animals treated with hyperthermia,
tumor accumulation of ELP1 (with Tb < Tt < Th) was
34% greater than that for ELP2 (for which
Tt > Th;
P < 0.005). From the combined data of the
two controls, we conclude that 60% of the total increase in ELP1
tumor accumulation versus the nonheated group can be
attributed to the effects of the ELP inverse transition, and 40% is
contributed by the physiological effects of hyperthermia. It is notable
that the percentage increase in tumor accumulation of the heated,
thermally responsive carrier versus both controls in the
radiolabel distribution experiments is consistent with the results from
the quantitative analysis of the window chamber experiments. Finally,
there was no significant difference in tumor localization between ELP1
and ELP2 for the unheated control group, indicating that the thermally
responsive and unresponsive carriers behaved similarly in the absence
of heat. This demonstrates that ELP2, with its composition and
molecular weight virtually identical to those of ELP1, is a
well-designed, thermally unresponsive control.
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2-fold increase in ELP1 accumulation in
heated tumors is not relative to free drug. Soluble
macromolecular carriers can provide up to several orders of magnitude
greater delivery and specificity (e.g., passive tumor
targeting due to the enhanced permeability and retention effect) to the
tumor versus free drug (1
, 6)
, particularly for
low molecular weight drugs with rapid renal clearance and for
hydrophobic drugs with low solubility. Therefore, the increase in tumor
delivery due to the ELP inverse temperature transition should be viewed
as a synergistic factor that further amplifies the advantages of using
macromolecular carriers to deliver drugs to solid tumors. Although other thermally responsive drug carriers such as temperature-sensitive liposomes and cross-linked hydrogels have been proposed (27, 28, 29) , to our knowledge, this is the first reported use of a thermally responsive polymer as a soluble carrier for thermal drug targeting. One advantage of this system over other thermally responsive carriers is that accumulation of the drug in the target tissue is driven through a phase transition of the carrier rather than through thermally triggered release of the drug. A concentration gradient is therefore not required to drive accumulation of the ELP, which will continue to accumulate through phase separation in a heated tumor (provided that Tt < Th) even when their blood concentration is less than the concentration in the tumor. This is an attractive feature because it enables the ELP-drug conjugate to be injected at a lower and therefore less toxic systemic concentration while still achieving a higher therapeutic concentration in the tumor.
Thermal targeting achieves regional targeting and is not specific to a particular cell type; therefore, any organ or tissue can be targeted independent of the availability of specific ligands or antibodies. For the delivery of radiotherapeutics, regional targeting (i.e., targeting to the tumor as opposed to tumor cells) is sufficient for therapy, and our initial work is therefore focused on their delivery to solid tumors. However, the delivery of other molecules such as chemotherapeutics, oligonucleotides, or imaging agents may require tumor cell-specific targeting. For these applications, thermal targeting could be synergistically combined with a cell-specific affinity-targeting moiety by gene level fusion or chemical conjugation of the targeting moiety (30, 31, 32) . Alternatively, the ELP could serve as a first stage, systemic targeting method to rapidly increase concentrations in the targeted region. This would be followed by release of the drug through labile linkers, which could also incorporate a secondary, affinity-targeting molecule.
The results presented here should be viewed as preliminary because no attempt has yet been made to optimize important experimental variables. In future studies, the ELP molecular weight is likely to be the most important design criterion to be explored because it is the primary design parameter that controls both plasma half-life and extravasation. The systemic concentration of the drug-ELP conjugate is a second important variable because of its effect on Tt, as described above, and because absolute systemic concentration may affect the relative distribution to the heated tumor versus other tissues. The administration protocol is the third variable that is likely to have a significant effect on the accumulation of thermally responsive ELPs in heated tumors. Controlled infusion, versus a bolus protocol as described in this study, may allow better control over the time course in vivo concentration to maintain the Tt in the requisite range (Tb < Tt < Th), thereby ensuring continual carrier accumulation in the tumor throughout the duration of the hyperthermia treatment. Furthermore, initial doses trapped in the hyperthermic region may trap later doses with higher efficiency. This may eventually allow a treatment protocol in which free ELP is initially injected, followed by a lower dose ELP conjugated to a therapeutic agent, thereby maximizing tumor-specific delivery of the therapeutic agent while minimizing systemic toxicity.
In summary, the results presented here demonstrate that a genetically
engineered, thermally responsive ELP in combination with hyperthermia
exhibits a 2-fold increase in accumulation in heated tumors compared
with the same polypeptide without hyperthermia. Comparison of these
results with the accumulation of a thermally unresponsive, control
polypeptide revealed that much of the increased accumulation of the
thermally responsive polypeptide in heated tumors was caused by the
inverse transition and subsequent aggregation of the thermally
responsive polypeptide rather than by the nonspecific, physiological
effects of hyperthermia. The window chamber studies also revealed that
the preferential accumulation of the thermally responsive polypeptide
in heated tumors is due to the combined effect of increased
accumulation in the vasculature and increased extravasation of the
polypeptide within the tumor. The results of this study are notable
because they demonstrate, for the first time to our knowledge, the
concept of thermal targeting through the exploitation of a
soluble-to-insoluble polymeric phase transition that has been designed
to occur at a specific temperature in vivo. We believe that
even greater accumulation of the thermally responsive carrier in tumors
may well be achieved after optimization of the this thermal targeting
system. Finally, synergistic thermal and affinity targeting schemes
offer the tantalizing possibility of further improving the targeted
delivery of therapeutics to solid tumors using these thermally
responsive polypeptides as macromolecular carriers.
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ACKNOWLEDGMENTS |
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FOOTNOTES |
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1 Supported by a grant from the Whitaker
Foundation (to A. C.), NIH Grant CA42745 (to M. W. D.), and
Department of Energy Grant DE-F602-96ER62148 (to M. R. Z.). 
2 To whom requests for reprints should be
addressed, at Department of Biomedical Engineering, Campus Box 90281,
Duke University, Durham, NC 27708-0281.
3 The abbreviations used are: ELP, elastin-like
polypeptide; SIB, N-succinimidyl 3-iodobenzoate;
Tb, physiological body temperature; Th,
temperature in the hyperthermic region; Tt, transition
temperature.
4 D. E. Meyer, and A. Chilkoti. Genetically
encoded synthesis of protein-based polymers with precisely specified
molecular weight by recursive directional ligation, manuscript in
preparation.
Received 10/30/00. Accepted 1/ 3/01.
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REFERENCES |
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TopABSTRACTINTRODUCTIONMATERIALS AND METHODSRESULTS AND DISCUSSIONREFERENCES |
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D. Raucher and A. Chilkoti Enhanced Uptake of a Thermally Responsive Polypeptide by Tumor Cells in Response to Its Hyperthermia-mediated Phase Transition Cancer Res., October 1, 2001; 61(19): 7163 - 7170. [Abstract] [Full Text] [PDF]
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) and in physiological saline (
).
B, percentage of free rhodamine as a function of
incubation time for the ELP-rhodamine conjugates in serum. 
, ELP2.
; ELP1 with heat,
) and
interstitium (