U.S. patent application number 16/141107 was filed with the patent office on 2019-03-28 for intracellular delivery.
This patent application is currently assigned to Massachusetts Institute of Technology. The applicant listed for this patent is Massachusetts Institute of Technology. Invention is credited to Andrea Adamo, Klavs F. Jensen, Robert S. Langer, Armon R Sharei.
Application Number | 20190093073 16/141107 |
Document ID | / |
Family ID | 48141314 |
Filed Date | 2019-03-28 |
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United States Patent
Application |
20190093073 |
Kind Code |
A1 |
Sharei; Armon R ; et
al. |
March 28, 2019 |
INTRACELLULAR DELIVERY
Abstract
A microfluidic system for causing perturbations in a cell
membrane, the system including a microfluidic channel defining a
lumen and being configured such that a cell suspended in a buffer
can pass therethrough, wherein the microfluidic channel includes a
cell-deforming constriction, wherein a diameter of the constriction
is a function of the diameter of the cell.
Inventors: |
Sharei; Armon R; (Cambridge,
MA) ; Adamo; Andrea; (Cambridge, MA) ; Langer;
Robert S.; (Newton, MA) ; Jensen; Klavs F.;
(Lexington, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Massachusetts Institute of Technology |
Cambridge |
MA |
US |
|
|
Assignee: |
Massachusetts Institute of
Technology
Cambridge
MA
|
Family ID: |
48141314 |
Appl. No.: |
16/141107 |
Filed: |
September 25, 2018 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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14352354 |
Apr 17, 2014 |
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PCT/US12/60646 |
Oct 17, 2012 |
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16141107 |
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61684301 |
Aug 17, 2012 |
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61548013 |
Oct 17, 2011 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
C12M 35/04 20130101;
C12M 35/02 20130101; B82Y 5/00 20130101; C12M 23/16 20130101; C12N
5/0602 20130101; C12N 15/87 20130101 |
International
Class: |
C12N 5/071 20100101
C12N005/071; C12M 1/42 20060101 C12M001/42; C12M 3/06 20060101
C12M003/06; C12N 15/87 20060101 C12N015/87 |
Goverment Interests
STATEMENT AS TO FEDERALLY-SPONSORED RESEARCH
[0002] This invention was made with Government support under Grant
No. RC1 EB011187 awarded by the National Institute of Health. The
Government has certain rights in the invention.
Claims
1. A microfluidic system for causing perturbations in a cell
membrane, the system comprising: a microfluidic channel defining a
lumen and being configured such that a cell suspended in a buffer
can pass therethrough, wherein the microfluidic channel includes a
cell-deforming constriction, wherein a diameter of the constriction
is a function of the diameter of the cell.
2. The microfluidic system of claim 1 wherein the diameter of the
constriction is substantially 20-99% of the diameter of the cell
passing therethrough.
3. The microfluidic system of claim 2, wherein a diameter of the
constriction is substantially 60% of the diameter of the cell.
4. The microfluidic system of claim 1, wherein the diameter of the
constriction is selected to induce temporary perturbations of the
cell wall large enough for a payload to pass through.
5. The microfluidic system of claim 4, wherein the diameter of the
constriction is also selected to reduce a likelihood that the cell
will die as a result of the deformation.
6. The microfluidic system of claim 1 wherein a cross-section of
the channel is selected from the group consisting of circular,
elliptical, an elongated slit, square, hexagonal, and
triangular.
7. The microfluidic system of claim 1 wherein the constriction
includes an entrance portion, a centerpoint, and an exit
portion.
8. The microfluidic system of claim 7 wherein the entrance portion
defines a constriction angle, wherein the constriction angle is
optimized to reduce clogging of the channel.
9. The microfluidic system of claim 7 wherein the entrance portion
defines a constriction angle, wherein the constriction angle is
optimized to improve delivery and cell viability.
10. The microfluidic system of claim 7 wherein the entrance portion
defines a 90 degree constriction angle.
11. The microfluidic system of claim 1 further comprising a
plurality of the microfluidic channels arranged in one of series
and parallel.
12. The microfluidic system of claim 1, further comprising a cell
driver.
13. The microfluidic system of claim 12, wherein the cell driver is
selected from a group consisting of: a pressure pump, a gas
cylinder, a compressor, a vacuum pump, a syringe, a syringe pump, a
peristaltic pump, a manual syringe, a pipette, a piston, a
capillary actor, a human heart, human muscle, and gravity.
14. The microfluidic system of claim 1, wherein a fluid flow of the
cell suspended in the buffer is channeled into the constriction,
the diameter of the constriction being greater than the diameter of
the cell passing therethrough, such that the cell is primarily
compressed by the fluid flow.
15-59. (canceled)
Description
RELATED APPLICATIONS
[0001] This application is a division of U.S. application Ser. No.
14/352,354, filed Apr. 17, 2014, entitled "INTRACELLULAR DELIVERY",
which is a national stage filing under 35 U.S.C. 371 of
International Patent Application Serial No. PCT/US2012/060646,
filed Oct. 17, 2012, which claims priority to U.S. Provisional
Application No. 61/548,013, filed Oct. 17, 2011, and U.S.
Provisional Application No. 61/684,301, filed Aug. 17, 2012, the
contents of each are hereby incorporated by reference.
BACKGROUND
[0003] Many pharmaceuticals largely focus on development of
small-molecule drugs. These drugs are so-called due to their
relatively small size that enables them to diffuse freely
throughout the body to reach their target. These molecules are also
capable of slipping across the otherwise impermeable cell membrane
largely unhindered. The next generation of protein, DNA or RNA
based therapies, however, cannot readily cross the cellular
membrane and thus require cellular modification to facilitate
delivery. Established methods use chemicals or electrical pulses to
breach the membrane and deliver the material into the cytoplasm.
Proper intracellular delivery is a critical step in the research,
development and implementation of the next generation of
therapeutics.
[0004] Existing methods are often difficult to develop and highly
specific to their particular application. Moreover, many clinically
important cell types, such as stem cells and immune cells, are not
properly addressed by existing methods. There is thus a need for
more robust and precise technique capable of addressing the needs
of modern biological/medical research.
SUMMARY
[0005] The invention is based on the surprising discovery that a
controlled injury, e.g., subjecting a cell to a constriction, rapid
stretching, rapid compression, or pulse of high shear rate, leads
to uptake of molecules into the cytoplasm of the cell from the
surrounding cell medium. Thus, the invention features a vector-free
microfluidic platform for direct-to-cytosol intracellular delivery
of materials, e.g., a compound or composition, to a eukaryotic
cell. The device is useful as a versatile and widely applicable
laboratory tool to deliver desired molecules into target cells. The
delivery of molecules into the cell using the methods described
herein is proportional, e.g., linearly or monotonically with cell
velocity through a constriction and/or pressure. For example, 50
.mu.l of cell suspension goes through the device in a few seconds.
The throughput ranges between 1 cell/second per channel (or even
less) to over 1,000 cells/second per channel. Typical cell
velocities through the constriction include 10 mm/second to 500
mm/second, although cell velocities can be up to 10 m/s (or even
higher). Additional channels can be placed in parallel to increase
the overall throughput of the system.
[0006] The uptake of molecule is diffusion-based rather than
endocytosis i.e., payload (compound(s) to be delivered to the cell)
are present in the cytoplasm rather than in endosomes following
passage through the device. Little or no payload appears in
endosomes following cell treatment. For example, large molecules
are taken up more slowly than smaller molecules. Controlled cell
stretching and velocity of movement of the cells through the
constriction leads to superior delivery of target molecules while
preserving the viability and integrity of the cells. After
treatment, cell viability is between 70-100%, e.g., typical
viability is 90% after treatment. By comparison, previous delivery
methods using high shear rates alone for seconds or milliseconds
have been shown to lead to poor viability of cells after treatment.
In contrast to prior techniques, the methods of the invention
subject the cells to a pulse of shearing ranging from 100-1000 Pa
for a very short period of time (approximately 100 microseconds) as
the cell passes through the constriction. The present techniques,
however, are fundamentally different from previous techniques. In
the present techniques, there is preferably an entire mechanical
deformation of the cell as it passes through the constriction,
which can impose different shearing forces than prior techniques.
In preferred embodiments, the cells are not subject to an electric
current. In other embodiments, a combination treatment is used,
e.g., mechanical deformation using the device described herein
followed by or preceded by electroporation (a type of osmotic
transfection in which an electric current is used to produce
temporary holes in cell membranes, allowing entry of nucleic acids
or macromoles).
[0007] A payload is a compound or composition to be delivered into
a cell. For example, a payload can include proteins, fluorescent
dies, quantum dots, carbon nanotubes, RNA molecules, DNA molecules,
antigens, and other macromolecules, nanoparticles, and compositions
of matter.
[0008] The width of the constriction of the device, the length of
the constricted portion, the geometry of the entrance region and
the channel depth of the device influence the delivery of molecules
into the cell. Preferably, the width of the constricted portion of
the conduit is no less than 4 .mu.m in diameter, and the length of
the constricted portion of the conduit is preferably between 40-50
.mu.m. The length of the constricted portion generally does not
exceed 90 .mu.m. The diameter of the constricted portion is related
to the type of cell to be treated. As is described below, the
diameter is less than the diameter of the cell (e.g., 20-99% of the
diameter of the cell). Many cells are between 5-15 .mu.m in
diameter, e.g. dendritic cells are 7-8 .mu.m in diameter. For
example, the diameter of the constriction portion is 4.5, 5, 5.5,
6, or 6.5 .mu.m for processing of single cells. In another example,
the size/diameter of the constricted portion for processing of a
human egg is between 6.2 .mu.m and 8.4 .mu.m, although larger and
smaller constrictions are possible (diameter of a human ovum is
approximately 12 .mu.m). In yet another example, embryos (e.g.,
clusters of 2-3 cells) are processed using a constriction diameter
of between 12 .mu.m and 17 .mu.m.
[0009] The device and methods are useful in vaccine development and
production using professional antigen presenting cells such as
dendritic cells. For example, a method of stimulating antigen
presentation is carried out by subjecting a dendritic cell to a
controlled injury such as transitory constriction or pulse of high
shear and contacting the dendritic cell with a solution comprising
a target antigen. The method yields highly activated antigen
presenting cells compared to previous methods of stimulation.
Vaccine production is carried out by propelling dendritic cells or
other antigen presenting cells through the constriction-containing
device (thereby subjecting the cells to a rapid stretching event)
and then incubating the cells in a solution containing the payload,
e.g., antigen. The cells are bathed in a cell culture medium
containing one or more antigens after rapid deformation of the
cells, but the cells may be contacted with the antigen prior to,
during, and/or after the rapid deformation event/process.
[0010] Surfactants (e.g., 0.1-10% w/w) are optionally used (e.g.,
poloxamer, animal derived serum, albumin protein) in the flow
buffer. Delivery of molecules into cells is not affected by the
presence of surfactants; however, surfactants are optionally used
to reduce clogging of the device during operation.
[0011] The device is made from silicon, metal (e.g., stainless
steel), plastic (e.g., polystyrene), ceramics, or any other
material suitable for etching micron scaled features and includes
one or more channels or conduits through which cells pass. Silicon
is particularly well suited, because micro patterning methods are
well established with this material, thus it is easier to fabricate
new devices, change designs, etc. Additionally, the stiffness of
silicon can provide advantages over more flexible substrates like
Polydimethylsiloxane (PDMS), e.g., higher delivery rates. For
example, the device includes 2, 10, 20, 25, 45, 50 75, 100 or more
channels. The device is microfabricated by etching the silicon.
Cells are moved, e.g., pushed, through the channels or conduits by
application of pressure. A cell driver can apply the pressure. A
cell driver can include, for example, a pressure pump, a gas
cylinder, a compressor, a vacuum pump, a syringe, a syringe pump, a
peristaltic pump, a manual syringe, a pipette, a piston, a
capillary actor, and gravity. As an alternative to channels, the
cells may be passed through a constriction in the form of a net or
closely-placed plates. In either case, the width of the
constriction through which the cells traverse is 20-99% of the
width or diameter of the cell to be treated in its natural, i.e.,
unstressed, state. Temperature can affect the uptake of
compositions and affect viability. The methods are carried out at
room temperature (e.g., 20.degree. C.), physiological temperature
(e.g., 39.degree. C.), higher than physiological temperature, or
reduced temperature (e.g., 4.degree. C.), or temperatures between
these exemplary temperatures.
[0012] Following controlled injury to the cell by constriction,
stretching, and/or a pulse of high shear rate, the cells are
incubated in a delivery solution that contains the compound or
molecule that one wishes to introduce into the cell. Controlled
injury may be characterized as small, e.g., 200 nm in diameter,
defect in the cell membrane. The recovery period for the cells is
on the order of a few minutes to close the injury caused by passing
through the constriction. The delivery period comprises 1-10
minutes or longer, e.g., 15, 20, 30, 60 minutes or more, with 2-5
minutes being optimal when operated at room temperature. Longer
time periods of incubation in the delivery solution do not
necessarily yield increased uptake. For example, the data indicated
that after 5 minutes, little or no additional material taken up by
the cells.
[0013] Thus, the invention provides a solution to long-standing
problems in the field of drug delivery to cells and to drawbacks
associated with earlier methods.
[0014] With respect to delivery of material to a eukaryote cell,
cells can be classified into two major categories:
[0015] 1) Easy-to-deliver (ETD) cells: Most available chemical and
viral methods fall under this category. Easy to deliver cells often
have no direct clinical relevance.
[0016] 2) Difficult-to-deliver (DTD) cells: High clinical
relevance. Advancements in delivery technology can greatly
enable/accelerate the development of novel therapies. This category
includes stem cells, primary cells, and immune cells. The market
for DTD delivery is expected to grow dramatically as novel RNA,
stem cell, and protein based therapeutics gain momentum in the
coming years.
[0017] The techniques described herein have proven especially
useful to DTD research areas, although the same techniques can be
used with ETD cells. In addition, it has facilitated the delivery
of materials (such as quantum dots, carbon nanotubes and
antibodies) that cannot be delivered effectively by any other
method to either ETD or DTD cells.
[0018] In general, in an aspect, implementations of the invention
can provide a microfluidic system for causing perturbations in a
cell membrane, the system including a microfluidic channel defining
a lumen and being configured such that a cell suspended in a buffer
can pass therethrough, wherein the microfluidic channel includes a
constriction, wherein a diameter of the constriction is a function
of the diameter of the cell.
[0019] Implementations of the invention may also provide one or
more of the following features. The diameter of the constriction is
substantially 20-99% of the diameter of the cell passing
therethrough. A cross-section of the channel is selected from the
group consisting of circular, elliptical, an elongated slit,
square, hexagonal, and triangular. The constriction includes an
entrance portion, a centerpoint, and an exit portion. The entrance
portion defines a constriction angle, wherein the constriction
angle is optimized to reduce clogging of the channel. The
microfluidic system further includes a plurality of the
microfluidic channels arranged in parallel, e.g., 2, 5, 10, 20, 40,
45, 50, 75, 100, 500, 1,000 or more.
[0020] In general, in another aspect, implementations of the
invention can also provide a method for delivering a compound into
a cell, the method including providing a cell in suspension or
suspending a cell and a payload in a solution, passing the solution
through a microfluidic channel that includes a constriction, sizing
the constriction as a function of the diameter of the cell, passing
the cell through the constriction such that a pressure is applied
to the cell causing perturbations of the cell large enough for the
payload to pass through, and incubating the cell in the solution
for a predetermined time after it passes through the
constriction.
[0021] Implementations of the invention may also provide one or
more of the following features. A diameter of the constriction is
substantially 20-99% of the diameter of the cell. A cross-section
of the microfluidic channel is selected from the group consisting
of circular, elliptical, an elongated slit, square, hexagonal, and
triangular. Passing the solution includes passing the solution
through an entrance portion, a centerpoint, and an exit portion of
the constriction. The method further includes reducing clogging of
the microfluidic channel by adjusting a constriction angle of the
entrance portion. The solution includes passing the solution
through a plurality of microfluidic channels arranged in
parallel.
[0022] In general, in still another aspect implementations of the
invention can also provide a method for delivering a compound into
a cell, the method including providing a cell in a solution or
suspending a cell in a solution, passing the solution through a
microfluidic channel that includes a constriction, sizing the
constriction as a function of the diameter of the cell, passing the
cell through the constriction such that a pressure is applied to
the cell causing perturbations of the cell, and incubating the cell
in the solution containing a payload for a predetermined time after
it passes through the constriction, wherein the perturbations are
large enough for the payload to pass through.
[0023] Implementations of the invention may also provide one or
more of the following features. A diameter of the constriction is
substantially 20-99% of the diameter of the cell. A cross-section
of the microfluidic channel is selected from the group consisting
of circular, elliptical, an elongated slit, square, hexagonal, and
triangular. Passing the solution includes passing the solution
through an entrance portion, a centerpoint, and an exit portion of
the constriction. The method further includes reducing clogging of
the microfluidic channel by adjusting a constriction angle of the
entrance portion. Passing the solution includes passing the
solution through a plurality of microfluidic channels arranged in
one of series and parallel. Incubating includes incubating the cell
for 0.0001 seconds to 20 minutes (or even longer). The pressure is
one of shearing and compression.
[0024] In general, in yet another aspect, implementations of the
invention can also provide a method for delivering a compound into
a cell, the method including providing a cell in a solution or
suspending a cell in a solution, deforming the cell such that
perturbations are caused in a membrane of the cell, and incubating
the cell in the solution with a payload after the cell has been
deformed.
[0025] Implementations of the invention may also provide one or
more of the following features. Deforming the cell includes
deforming the cell for 1 .mu.s to 10 ms, e.g., 10 .mu.s, 50 .mu.s,
100 .mu.s, 500 .mu.s, and 750 .mu.s. Incubating occurs for 0.0001
seconds to 20 minutes, e.g., 1 second, 30 seconds, 90 seconds, 270
seconds, and 900 seconds.
[0026] Various implementations of the invention may provide one or
more of the following capabilities. Greater precision and
scalability of delivery can be achieved when compared with prior
techniques. Delivery of a material to a cell can be automated.
Material such as proteins, RNA, siRNA, peptides, DNA, and
impermeable dye can be implanted into a cell, such as embryonic
stem cells or induced pluripotent stem cells (iPSCs), primary cells
or immortalized cell lines. The device and methods are amenable to
any cell type, and the size of the constricted portion is tailored
to the of the cell to be treated. The devices and methods can
provide significant advantages. For example, experimental noise in
current systems can be reduced when compared with prior techniques.
Delivery quantities of a material can be consistent across the cell
population. Cells can be individually handled rather than being
handled as a batch. The invention has also demonstrated a fairly
unique opportunity to deliver a variety of nanoparticles and
proteins to the cytosol. Existing methods are fairly unreliable or
inefficient at performing such functions.
[0027] With respect to delivery of sensitive payloads, e.g.,
proteins (especially large proteins, e.g., greater than 30, 50,
100, 150, 200, 300, 400, 500 kDa or more), quantum dots, or other
payloads that are sensitive to or damaged by exposure to
electricity, are reliably delivered into cells while preserving the
integrity and activity of the sensitive payload. Thus, the device
and methods have significant advantages over existing techniques
such as electroporation, which subjects payload compositions to
electricity (thereby damaging the payload) and leads to low cell
viability (e.g., 505 or more of the cells typically die after
electroporation). Another advantage of the rapid
stretch/deformation method is that stem or precursor cells are
rendered receptive to uptake of payload without altering the state
of differentiation or activity of the treated cell. In addition to
delivery of compositions into the cytoplasm of the cell for
therapeutic purposes, e.g., vaccine production, the method is used
to introduce molecules, e.g., large molecules comprising a
detectable marker, to label intracellular structures such as
organelles or to label intracellular constituents for diagnostic or
imaging purposes.
[0028] Various implementations of the invention may also provide
one or more of the following capabilities. DNA can be delivered
into dose-to-deliver cells such as stem, primary, immune cells.
Delivery of very large plasmids (even entire chromosomes) can be
accomplished. Quantitative delivery into cells of known amount of a
gene construct to study the expression level of a gene of interest
and its sensitivity to concentration can also readily be
accomplished. Delivery of known amounts of DNA sequences together
with known amount of enzymes that enhance DNA recombination in
order to achieve easier/more efficient stable delivery, homologous
recombination, and site-specific mutagenesis can be accomplished.
The methods and devices described herein can also be useful for
quantitative delivery of RNA for more efficient/conclusive RNA
studies. Delivery of small interfering RNA (siRNA) into the
cytoplasm of a cell is also readily accomplished.
[0029] Various implementations of the invention may also provide
one or more of the following capabilities. RNA can be delivered
into a cell for RNA silencing without the need for liposomes. Known
amounts of RNA molecules together with known amounts of dicer
molecules can be delivered to achieve standardized, efficient, RNA
across multiple cell lines in different conditions. mRNA can be
delivered into cells to study aspects of gene expression
regulations at the posttranscriptional level. Known amounts of
label of RNA to study the half-life of RNAs and cells can be
possible. Universal protein delivery can be achieved. Known amounts
of label proteins can be delivered to study their half-life in
cells. Delivery of label proteins to study protein localization can
be accomplished. Known amounts of tagged proteins can be delivered
to study protein-protein interactions in the cellular environment.
Delivery of labeled antibodies into living cells for immunostaining
and fluorescence-based Western blotting can be achieved.
[0030] Various implementations of the invention may also provide
one or more of the following clinical and research capabilities.
Quantitative delivery of drugs to cell models for improved
screening and dosage studies can be achieved. The method could be
deployed as a high throughput method of screening protein activity
in the cytosol to help identify protein therapeutics or understand
disease mechanisms. Such applications are presently severely
limited by current protein delivery methods due to their
inefficiencies. The devices and techniques are useful for
intracellular delivery of drugs to a specific subset of circulating
blood cells (e.g. lymphocytes), high throughput delivery of sugars
into cells to improve cryopreservation of cells, especially
oocytes, targeted cell differentiation by introducing proteins,
mRNA, DNA and/or growth factors, delivery of genetic or protein
material to induce cell reprogramming to produce iPS cells,
delivery of DNA and/or recombination enzymes into embryonic stem
cells for the development of transgenic stem cell lines, delivery
of DNA and/or recombination enzymes into zygotes for the
development of transgenic organisms, DC cell activation, iPSC
generation, and stem cell differentiation, nano particle delivery
for diagnostics and/or mechanic studies as well as introduction of
quantum dots. Skin cells used in connection with plastic surgery
are also modified using the devices and method described
herein.
[0031] A method of stimulating antigen presentation using the
method to deliver antigen and/or immune stimulatory molecules
yields antigen presenting cells, e.g., dendritic cells, with
improved levels of activity compared to convention methods of
stimulation, thereby leading to increased levels of T and B-cell
mediated immunity to a target antigen. Such a method could thus be
employed as a means of activating the immune system in response to
cancer or infections
[0032] For screening, imaging, or diagnostic purposes, the device
is used in a method of labeling cells. A method of labeling a cell
is carried out by subjecting a cell to a controlled injury and
contacting the cell with a solution comprising a detectable marker,
wherein said injury comprises a transitory constriction or pulse of
high shear. The detectable marker comprises a fluorescent molecule,
a radionuclide, quantum dots, gold nanoparticles, or magnetic
beads.
[0033] Prior to the invention, manipulation of stem cells for the
purpose of introducing exogenous compositions has been difficult.
The device and methods described herein, e.g., passage of stem
cells or progenitor cells such as induced pluripotent stem cells
(iPSCs) through a constriction channel does not induce
differentiation, but does reliably induce uptake of compositions
into the cell. For example, differentiation factors are introduced
into such cells. After uptake of introduced factors, the cells
proceed on a differentiation pathway dictated by the introduced
factor without complications associated with the method by which
the factor(s) was introduced into the cell.
[0034] In addition to single cells, even very large cells, e.g.,
eggs; approximately 200 .mu.m in diameter, clusters of cells, e.g.,
2-5 cell clusters such as an embryo comprising 2-3 cells, are
treated to take up target compositions. The size of the aperture is
adjusted accordingly, i.e., such that the width of the constriction
is just below the size of the cluster. For example, the width of
the channel is 20-99% of the width of the cell cluster.
[0035] Cells or cell clusters are purified/isolated or enriched for
the desired cell type. Dendritic cells or other cells, e.g., immune
cells such as macrophages, B cells, T cells, or stem cells such as
embryonic stem cells or iPS, used in the methods are purified or
enriched. For example, cells are isolated or enriched by virtue of
their expression of cell surface markers or other identifying
characteristics. Dendritic cells are identified and isolated by
virtue of their expression of the 0-intergrin, CD11c or other
identifying cell surface markers. With regard to cells, the term
"isolated" means that the cell is substantially free of other cell
types or cellular material with which it naturally occurs. For
example, a sample of cells of a particular tissue type or phenotype
is "substantially pure" when it is at least 60% of the cell
population. Preferably, the preparation is at least 75%, more
preferably at least 90%, and most preferably at least 99% or 100%,
of the cell population. Purity is measured by any appropriate
standard method, for example, by fluorescence-activated cell
sorting (FACS).
[0036] Payload compositions such as polynucleotides, polypeptides,
or other agents are purified and/or isolated. Specifically, as used
herein, an "isolated" or "purified" nucleic acid molecule,
polynucleotide, polypeptide, or protein, is substantially free of
other cellular material, or culture medium when produced by
recombinant techniques, or chemical precursors or other chemicals
when chemically synthesized. Purified compounds are at least 60% by
weight (dry weight) the compound of interest. Preferably, the
preparation is at least 75%, more preferably at least 90%, and most
preferably at least 99%, by weight the compound of interest. For
example, a purified compound is one that is at least 90%, 91%, 92%,
93%, 94%, 95%, 98%, 99%, or 100% (w/w) of the desired compound by
weight. Purity is measured by any appropriate standard method, for
example, by column chromatography, thin layer chromatography, or
high-performance liquid chromatography (HPLC) analysis. A purified
or isolated polynucleotide (ribonucleic acid (RNA) or
deoxyribonucleic acid (DNA)) is free of the genes or sequences that
flank it in its naturally-occurring state. Examples of a an
isolated or purified nucleic acid molecule include: (a) a DNA which
is part of a naturally occurring genomic DNA molecule, but is not
flanked by both of the nucleic acid sequences that flank that part
of the molecule in the genome of the organism in which it naturally
occurs; (b) a nucleic acid incorporated into a vector or into the
genomic DNA of a prokaryote or eukaryote in a manner, such that the
resulting molecule is not identical to any naturally occurring
vector or genomic DNA; (c) a separate molecule such as a cDNA, a
genomic fragment, a fragment produced by polymerase chain reaction
(PCR), or a restriction fragment; and (d) a recombinant nucleotide
sequence that is part of a hybrid gene, i.e., a gene encoding a
fusion protein. Isolated nucleic acid molecules according to the
present invention further include molecules produced synthetically,
as well as any nucleic acids that have been altered chemically
and/or that have modified backbones.
[0037] A suspension solution is any physiologic or cell-compatible
buffer or solution. For example, a suspension solution is cell
culture media or phosphate-buffered saline. A payload is the same
or different suspension solution, which also contains the
composition intended to be delivered inside the cell.
[0038] Advantages of the device include avoiding modification of
the desired payload, and not necessarily exposing the payload to
any electromagnetic fields or other forms of stress. With respect
to electroporation, this method has been shown to damage proteins
and be ineffective in delivery. This significant drawback is not an
issue with the method described herein; the present method is
particularly suitable for delivery of sensitive payloads, e.g.,
proteins, particularly large proteins (e.g., 40 kDa-70 kDa, and up
to 120, 130, 150, 200 kDa or more), large nucleic acid constructs
(e.g., plasmids and other constructs containing 1 kb, 2 kb, 5 kb,
or more of nucleic acid polymers and up to entire chromosomes),
large compounds, as well as quantum dots (e.g., 12 nm in diameter)
and other materials that are known to be sensitive and easily
damaged upon exposure to electricity. For example, the surface
ligands on a nanoparticle or quantum dot can be damaged or become
charged in response to an electric field thus resulting in
aggregation of the particles thereby limiting/eliminating their
functionality. Yet another advantage of the controlled injury
method is the timing of contacting the cells with the delivery
composition. Particularly relevant for proteins, which are
sensitive to proteases, temperature, as well as electricity, cells
are contacted with payload solution after treatment and for a
relatively short period of time compared to earlier methods. The
microfluidic nature of the device also requires far smaller working
volumes thereby conserving precious raw materials and/or cells. The
device can also be coupled with existing delivery methods such as
electroporation or liposomes to produce a greatly enhanced delivery
relative to each method individually.
[0039] Functional activity of delivered payload is inversely
correlated to fluid shear stress, i.e., physical strain to the cell
membrane such as stretching of the cell membrane mediates uptake of
payload rather than shear. Conventional nanoparticle delivery
methods may result in greater amounts of material gaining access to
the intracellular environment of the cell; however, those methods
lead to less activity of the delivered material compared to the
methods described herein due to the fact that previous methods
result in sequestration of the delivered material in endosomes. The
methods described herein lead to direct-to-cytosol delivery of
compounds/compositions such that a lesser amount of payload
delivered into the cell leads to a greater amount of functional
activity of the delivered molecules due to their accessibility to
other cytosolic components. For example, earlier methods for
delivering nanoparticles have resulted in 2-10 times the amount of
delivered material into the cell but with little or no functional
activity of the delivered material due to sequestration in
endosomes. The devices and methods of the invention overcome this
drawback of previous intracellular delivery methods by avoiding the
endosomal compartment.
[0040] Additional advantages and features include time scale of
treatment and cell speeds that are much faster than earlier
approaches. Moreover, other methods do not squeeze the cells as
hard as the present methods, e.g., as determined by size (diameter)
of cell relative to size (diameter) of constriction (as a % of the
diameter of the cell). This rapid, forceful, but sub-leather,
squeeze or deformation leads to superior results in
direct-to-cytosol payload uptake by cells. Deformation of the cell
is sudden, i.e., occurs over substantially 1 .mu.s to 1 ms. In
general, too much deformation induced cell stress can be lethal to
the cell, while at the same time, too little stress does not induce
cell perturbations. Therefore the current subject matter provides
methods and systems that cause sufficient stress to induce
temporary perturbations but not so much stress that the
perturbations are permanent and lethal to the cell.
[0041] Any of the methods described above are carried out in vitro,
ex vivo, or in vivo. For in vivo applications, the device may be
implanted in a vascular lumen, e.g., an in-line stent. These and
other capabilities of the invention, along with the invention
itself, will be more fully understood after a review of the
following figures, detailed description, and claims.
BRIEF DESCRIPTION OF THE FIGURES
[0042] FIG. 1a is a schematic diagram of a microfluidic system.
Cells are exposed to the delivery material (payload) after passing
through the constriction.
[0043] FIG. 1b is a schematic diagram of a microfluidic system.
Cells are exposed to the delivery material (payload) throughout the
process by suspending the cells in a solution that includes the
delivery material (payload) (e.g., the cells are exposed to the
delivery material before and after passing through the
constriction).
[0044] FIG. 2A is a schematic diagram of an embodiment of a
microfluidic system.
[0045] FIG. 2B is an illustration diagram of a microfluidic system
depicting depth, width, and length.
[0046] FIG. 3 is a schematic diagram of a microfluidic system.
[0047] FIG. 4 is a schematic diagram showing perturbations in a
cell wall.
[0048] FIG. 5 is a photograph of a microfluidic system.
[0049] FIG. 6 is a photograph of a microfluidic system.
[0050] FIG. 7 is a photograph of a microfluidic system.
[0051] FIGS. 8a-8b are graphs showing exemplary results obtained
from a microfluidic system.
[0052] FIG. 9 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0053] FIG. 10 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0054] FIG. 11 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0055] FIG. 12 is a schematic diagram of a microfluidic system.
[0056] FIG. 13 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0057] FIG. 14 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0058] FIG. 15 is a graph showing exemplary results obtained from
cells that were processed using a microfluidic system.
[0059] FIGS. 16a-16f are exemplary schematic diagrams of
microfluidic systems.
[0060] FIG. 17 is a flow diagram relating to a method of using a
microfluidic system.
[0061] FIGS. 18a-18b are graphs showing exemplary results obtained
from cells that were processed using a microfluidic system.
[0062] FIG. 19 is an overlay of transmission and confocal
fluorescence images, followed by z-section confocal fluorescence
images of treated cells delivered with quantum dots (QDs) using the
current subject matter.
[0063] FIG. 20A illustrates delivery efficiency into HeLa cell
cytosol upon current subject matter treatment with QDs coated with
poly-imidazole ligand (PIL). Cell viability was >80% as measured
by flow cytometry.
[0064] FIG. 20B illustrates viability of HeLa cells upon delivery
of plain QD535 by the current subject matter, as measured by
propidium iodide staining and flow cytometry measurement.
[0065] FIG. 21 illustrates construct design, absorbance, and
stability in various media.
[0066] FIG. 22A illustrates live cell confocal microscopy images of
treated and control cells.
[0067] FIG. 22B illustrates a change in intensity of treated cells
as a function of time in the green and red channels.
[0068] FIGS. 23A-23D illustrate flow cytometry measurements of
average cell fluorescence and viability.
[0069] FIG. 24 illustrates epifluorescence imaging of unaggregated
single quantum dots within the cell cytosol after device treatment
with a 10 nM quantum dot solution, and blinking traces of three
quantum dots with autofluorescence.
[0070] FIG. 25 illustrates experimental results showing that
delivery performance depends on cell speed and constriction
design.
[0071] FIG. 26 illustrates scans of different horizontal planes of
a HeLa cell after the delivery of pacific blue conjugated 3 kDa
dextran, as measured by confocal microscopy.
[0072] FIG. 27 illustrates a simplified, 2D diffusion model that
simulates passive diffusion of material into a cell across a
porated membrane.
[0073] FIG. 28 illustrates the results of a two-tiered delivery of
material.
[0074] FIG. 29 illustrates data relating to SiRNA, protein, and
nanoparticle delivery.
[0075] FIG. 30 illustrates applicability of the current subject
matter across cell types.
[0076] FIG. 31 illustrates data from nanomaterial and antibody
delivery.
[0077] FIG. 32 illustrates protein delivery applications.
[0078] FIG. 33 is a table of exemplary cell types, which payload
has successfully been delivered.
[0079] FIG. 34 is an illustration depicting a system in which a
patient's blood is treated by a microfluidic device for the
delivery of payload such as macromolecules.
[0080] FIG. 35 illustrates delivery efficiency and viability of
human embryonic stem cells treated with a 10 .mu.m-6 .mu.m device
to deliver payload.
[0081] FIGS. 36A-36F depict the generation and characterization of
mouse and human iPSC lines by direct delivery of fused
reprogramming proteins using the current subject matter.
[0082] FIG. 37 depicts preliminary protein reprogramming results
and depicts expression of the human embryonic stem cell marker
Oct4, SSEA-4, Tra-60, Tra-80, Alkaline Phosphatase (AP) in iPSC
colonies.
[0083] FIG. 38 depicts micrographs illustrating a device modified
by incorporated electrodes on either side of the constriction by
photolithographic patterning and Au deposition to introduce a
localized electrical field into the channel thereby combining cell
deformation with electroporation.
[0084] FIG. 39 depicts another embodiment of the microfluidic
system wherein entrance portion has a constriction angle of 90
degrees.
[0085] FIGS. 40A and 40B are plots showing a comparison of
viability and delivery efficiency between a device in accordance
with the example embodiment depicted in FIG. 2A and a device in
accordance with an example embodiment depicted in FIG. 39.
[0086] FIG. 41 is a histogram of CD45 expression of activated T
cells as measured by an Alexa 488 antibody to CD45. Cells that are
treated by the device in the presence of CD45 silencing RNA exhibit
a lower fluorescence intensity peak thereby indicating knockdown of
CD45 gene expression.
[0087] FIG. 42 is an illustration depicting several example fields
of application such as regenerative medicine; immunology; imaging
and sensing; and cancer vaccines and cancer research.
[0088] FIGS. 43A and 43B are intensity histograms from flow
cytometry of a control population that is exposed to cascade blue
conjugated 3 kDa dextran and a population of cells that have been
subjected to a 30 .mu.m-6 .mu.m device and then exposed to the 3
kDa dextran.
[0089] FIG. 44 is a bar graph illustrating GFP knockdown in human
embryotic stem cells after treatment using the microfluidic device
and related methods.
[0090] FIGS. 45A and 45B are two plots illustrating the dye
intensity and viability of human embryotic stems cells after
delivery of a 3 kDa blue dye.
DETAILED DESCRIPTION
[0091] Embodiments of the invention provide techniques for applying
controlled deformation to a cell for a predetermined amount of time
in order to cause perturbations in the cell membrane such that
materials can be delivered to the inside of the cell. The
deformation can be caused by, for example, pressure induced by
mechanical strain or shear forces. In one example, a microfluidic
system includes a structure that controls and/or manipulates fluids
by geometrically confining the fluids on a small scale (eg., sub
milliliter volumes such as microlitres, nanoliters, or picoliters).
The microfluidic system is capable of intracellularly delivering
virtually any payload into a cell. The system consists of one or
more microfluidic channels with a constriction that the cells pass
through. Preferably, the cells flow through the microfluidic
channel suspended in a liquid medium that is pressure driven
through the system. When a cell passes through the constriction,
its membrane is perturbed causing temporary disruptions in the
membrane and resulting in the uptake of the payload that is present
in the surrounding media. The constriction is a function of the
size of the target cell, but preferably on the same order or
smaller than the cell diameter. Multiple constrictions can be
placed in parallel and/or series. The perturbation in the cell is a
breach in the cell that allows material from outside the cell to
move into the cell (e.g., a hole, tear, cavity, aperture, pore,
break, gap, perforation). The perturbations (e.g., pores or holes)
created by the methods described herein are not formed as a result
of assembly of protein subunits to form a multimeric pore structure
such as that created by complement or bacterial hemolysins. Other
embodiments are within the scope of the described subject
matter.
[0092] Referring to FIGS. 1-3, a microfluidic system 5 includes a
channel 10 defining a tubular lumen. The microfluidic channel 10
includes a constriction 15 that is preferably configured such that
only a single target cell 20 can pass through the constriction 15
at one time. Preferably, the cells 20 pass through the channel 10
suspended in a solution buffer 25 that also includes delivery
materials 30, although the delivery materials can be added to the
solution buffer 25 after the cells 20 pass through the constriction
15. As the cell 20 approaches and passes through the constriction
15, the constriction 15 applies pressure (e.g., mechanical
compression) to the cell 20, squeezing the cell 20 (e.g., shown as
cell 20.sub.1). The pressure applied to the cell by the
constriction 15 causes perturbations (e.g., holes shown in FIG. 4)
in the cell membrane (e.g., cell 202). Once the cell passes through
the constriction 15, the cell 20 begins to uptake the material in
the solution buffer 25 through the holes, including the delivery
material 30 (e.g., cell 203). The cell membrane recovers over time,
and at least a portion of the delivery material 30 preferably
remains trapped inside the cell.
[0093] The configuration of the constriction 15 can be customized
to control the constriction of the cell 20, thereby controlling the
pressure applied to the cell 20. Preferably, the constriction 15
includes an entrance portion 35, a centerpoint 40, and exit portion
45. For example, the diameter(s) of the constriction 15 can be
varied to adjust the pressure applied to the cell (and how quickly
that pressure is applied/released), and the length of the
constriction 15 can be varied to adjust the amount of time pressure
is applied to the cell. In certain configurations, physical
constriction of the cell is not required, rather very briefly
subjecting the cell to an unusually high sheer rate and/or
compression rate may cause the desired perturbations. Generally,
there is no requirement relating to the outside diameter of the
microfluidic system and the ratio of the inner diameter to the
outer diameter can be varied (e.g., greater than 5).
[0094] The diameter of the centerpoint 40 can be a function of the
diameter of the cell 20. Preferably, the centerpoint 40 is on the
same order as or smaller than the diameter of the cell 20 (e.g.,
20-99% of the diameter of the cell). Preferably, the diameter of
the centerpoint 40 is between 60% and 70% of the diameter of the
cell, although optimal centerpoint diameter can vary based on the
application and/or cell type. Exemplary diameters of the
centerpoint 40 that has been used in prior experiments is 5-6
.mu.m, and 7-8 .mu.m. The centerpoint 40 can also be larger than
the diameter of the cell 20, but be configured to cause a pulse of
pressure (e.g., shearing) that is applied to the cell 20. Such
pressure can be applied to the cell 20 without it touching the
walls of the channel 10. Shear can be measured by known techniques
(e.g, Journal of Applied Physics 27, 1097 (1956); Murphey et
al.).
[0095] The constriction angle (e.g., .alpha. in FIG. 2A) of the
entrance portion 35 can vary (e.g., how quickly the diameter
decreases). The constriction angle is preferably an angle that
minimizes clogging of the system 5 while cells are passing
therethrough. The angle of the exit portion 45 can vary as well.
For example, the angle of the exit portion 45 is configured to
reduce the likelihood of turbulence/eddies that can result in
non-laminar flow (e.g., a range from 1-80 degrees). The walls of
the entrance portion 35 and/or the exit portion 45 are preferably
linear, although other configurations are possible (e.g., the walls
can be curved).
[0096] The cross-section of the channel 10, the entrance portion
35, the centerpoint 40, and the exit portion 45 can vary. For
example, the various cross-sections can be circular, elliptical, an
elongated slit, square, hexagonal, triangular, etc. The length of
the centerpoint 40 can also vary, and can be adjusted to vary the
amount of time that pressure is applied to the cell 20 as it passes
through the constriction 15. At a given flow rate, a longer
constriction 15 (e.g., a longer centerpoint 40) will apply pressure
to cell 20 for a longer period of time. The depth of the channel
10, the entrance portion 35, the centerpoint 40, and the exit
portion 45 can vary. For example, the depth can be adjusted to
provide a tighter constriction and thereby enhance delivery in a
manner similar to changes in constriction width. Width and lenge
vary between device designs and can be determined during
manufacture of the device, such as by a chrome mask used in a
lithography step (when the device is silicon based). Depth can be
uniform throughout the channel and can be determined during
manufacture of the device, such as by a deep reactive ion etching
step. The depth can be, for example, 15 .mu.m-20 .mu.m. As used
herein, device dementions are denoted by a series of numbers
indicating length, width, and number of constrictions (e.g., 30
.mu.m-6 m.times.5 denotes a device with a 30 .mu.m length, 6 .mu.m
width, and 5 constrictions).
[0097] The velocity at which the cells 20 pass through the channel
10 can also be varied to control delivery of the delivery material
30 to the cells 20. For example, adjusting the velocity of the
cells 20 through the channel 10 can vary the amount of time that
pressure is applied to the cells, and can vary how rapidly the
pressure is applied to the cell (e.g., slowly or shockingly). The
cells 20 pass through the system 5 at a rate of at least 0.1 mm/s
such as 0.1 mm/s to 5 m/s, and preferably between 10 mm/s to 500
mm/s, although other speeds are possible. In some embodiments, the
cells 20 can pass through the system 5 at a rate greater than 5
m/s.
[0098] The channel 10 can be fabricated from various materials such
as silicon, glass, ceramics, crystalline substrates, amorphous
substrates, and polymers (e.g., Poly-methyl methacrylate (PMMA),
PDMS, Cyclic Olefin Copolymer (COC), etc). Fabrication is
preferably clean-room based, and can use, for example dry etching,
wet etching, photolithography, injection molding, laser ablation,
SU-8 masks, etc. One exemplary channel 10 is approximately 40-50
.mu.m long, having a non-constriction diameter of approximately 50
.mu.m, having a constriction diameter of approximately 4-8 .mu.m.
Preferably, the length of the channel 10 is kept as short to avoid
clogs. Other dimensions are possible.
[0099] FIG. 39 depicts another embodiment of the microfluidic
system. In this embodiment, channel 10 includes a preliminary
entrance portion 50 that does not constrict the cell 20. An
expanded channel portion 55 provides for entrance portion 35 to
have a constriction angle of 90 degrees (e.g., alpha in FIG.
2A).
[0100] FIGS. 40A and 40B are two plots showing a comparison of
viability and delivery efficiency between two example embodiments.
Label 4000 designates measurements taken while using an embodiment
in accordance with FIG. 2A while 4010 designates measurements taken
while using an embodiment in accordance with FIG. 39. For the same
cell speed and operating pressure, the embodiment of FIG. 39 has
been shown to have high delivery efficiency and viability. This is
despite having similar shear rates, cell speed, and time spent
under compression as the embodiment of FIG. 2A.
[0101] Several parameters can influence the delivery of the
delivery material 30 into the cell 20. For example, the dimensions
of the constriction 15, the operating flow speeds (e.g., cell
transit time to the constriction 15), concentration of the delivery
material 30 in the solution buffer 25, and the amount of time that
the cell 20 recovers/incubates in the solution buffer 25 after
constriction can affect the absorption of the delivery material 30
into the cell 20. Additional parameters influencing the delivery of
the material 30 into the cell 20 can include the velocity of the
cell 20 in the constriction 15, the shear rate in the constriction
20, the velocity component that is perpendicular to flow velocity,
a cell compression rate, and time in constriction. Such parameters
can be designed to control delivery of the delivery material 30.
The composition of the solution buffer 25 (e.g., salt
concentration, serum content, etc.) can also impact delivery of the
delivery material 30. As the cell 20 passes through the
constriction 15, the deformation/stress induced by the constriction
15 temporarily causes injury to the cell that causes passive
diffusion of material through the perturbation. In some
embodiments, the cell 20 is only deformed for brief period of time,
on the order of 100 .mu.s to minimize the chance of activating
apoptotic pathways through cell signaling mechanisms, although
other durations are possible (e.g., ranging from nanoseconds to
hours). Initial observations have indicated that absorption of the
delivery material 30 by the cell 20 occurs on the order of minutes
after the cell 20 passes through the constriction 15.
[0102] The cells 20 can be driven through the channel 10 by various
methods. For example, pressure can be applied by a pump on the
entrance side (e.g., gas cylinder, or compressor), a vacuum can be
applied by a vacuum pump on the exit side, capillary action through
a tube, and/or the system 5 can be gravity fed. Displacement based
flow systems can also be used (e.g., syringe pump, peristaltic
pump, manual syringe or pipette, pistons, etc.). Exemplary flow
rates through a single channel 10 are on the order of 1 .mu.l in a
few seconds. Additionally, solution buffer 25 can include one or
more lubricants (pluronics or other surfactants) that can be
designed to reduce or eliminate clogging of the channel 10 and
improve viability.
[0103] The system 5 can be controlled to ensure that delivery
quantities of the delivery material 30 is consistent across the
cell population. For example, the system 5 can include the use of a
post-constriction convective delivery mechanism that impinges
delivery material 30 onto the permeabalized cell membrane of the
cell 20. By controlling the flow rate of the secondary stream, the
quantity of delivery material 30 provided to the cell can
preferably be controlled. Additionally, controlling the
concentration of delivery material 30 in the solution buffer 25
during membrane recovery can also improve the consistency of
delivery of the delivery material 30 to the population of cells.
Preferably, the system 5 operates as a purely mechanical system
without applying any electrical fields and/or chemical agents,
although other configurations are possible (e.g., electrical and/or
optical sensors can be used to measure cell properties such as
fluorescence). Additionally, the system 5 preferably operates
independent of the type of material being delivered. For example,
proteins, RNA, and DNA can be delivered through the same system
without any additional modifications.
[0104] In some configurations with certain types of cells 20, the
cells 20 can be incubated in one or more solutions that aid in the
absorption of the delivery material to the interior of the cell.
For example, the cells 20 can be incubated in a depolymerization
solution such as Lantrunculin A (0.1 .mu.g/ml) for 1 hour prior to
delivery to depolymerize the actin cytoskeleton. As an additional
example, the cells can be incubated in 10 .mu.M Colchicine (Sigma)
for 2 hours prior to delivery to depolymerize the microtubule
network. These methods can help in obtaining gene expression when
delivering DNA.
[0105] Referring also to FIG. 5, a photograph of a parallel
configuration of the system 5 is shown. The system 5 can include
any number of parallel channels. Preferably, as additional parallel
channels are added to the system 5, the overall throughput of the
system 5 can be increased. FIG. 6 shows a photograph of a parallel
configuration of the system 5 that includes filters at the inlet of
each of the channels 10. Additionally, FIG. 6 also shows a
configuration of the constriction 15 that includes an entrance
portion 35 that includes multiple steps. Referring also to FIG. 7,
an additional photograph of a prototype of the system 5 is shown.
As evident in FIG. 7, the prototype, including incubation well, has
dimensions of approximately 1 inch.times.1/4 inch.times.1/4 inch.
Other configurations of the system 5 can also include sorters,
pretreatment/post treatment modules, and/or sensor modules (e.g.,
optical, electrical, and magnetic).
[0106] As described in more detail below with regard to the
examples, the microfluidic system and related methods have a broad
range of applications. FIG. 42 is an illustration depicting several
example fields of application. For example, the current subject
matter can be applied to regenerative medicine such as to enable
cell reprogramming and stem cell differentiation. The current
subject matter can be applied to immunology such as for antigen
presentation and enhancement/suppression of immune activity through
delivery to dendritic cells, monocytes, T cells, B cells and other
lymphocytes. Further, imaging and sensing can benefit from improved
delivery to target cells of quantum dots, carbon nanotubes and
antibodies. Additionally, the current subject matter has
application in cancer vaccines and research, such as for
circulating tumor cell (CTC) isolation and Lymphoma treatment. The
method also provides a robust platform to screen for active siRNA
and small molecule compounds capable of treating a disease or
manipulating cell behavior.
[0107] This concept has been successfully demonstrated in a
prototype where the cells 20 were induced to take-up otherwise
membrane-impermeable dye (e.g., fluorescent dyes from 3 kDA to 2
MDA in molecular mass, DNA, protein, RNA, nanotubes or
nanoparticles present in the solution buffer 25. The cells 20 have
been shown to recover and proliferate after the process while
retaining the delivered material for over 72 hours. Eleven
different cell types have been tested with this system, including
those listed in FIG. 33, hence demonstrating that the system
provides robust performance in different cell types. FIG. 33 is a
table including cell types which the current subject matter has
successfully been applied. Average cell throughput has been
measured on the order of 5,000-20,000 cells/second, average
delivery efficiency has been measured at 96%, and cell viability
has been measured at 95% using a single channel 10. All tests were
performed at room temperature. Temperature, however, may be varied
in some techniques. For example, the methods can be carried out at
room temperature (e.g., 20.degree. C.), physiological temperature
(e.g., 39.degree. C.), higher than physiological temperature, or
reduced temperature (e.g., 4.degree. C.), or temperatures between
these exemplary temperatures. Performing the methods at a reduced
temperature (i.e. substantially near 4.degree. C. which can be
achieved, for example, by using refrigeration, ice bath, or other
known techniques), has produced a surprising improvement in
delivery efficiency and cell viability. Thus, the temperature can
be adjusted to affect composition delivery and cell viability.
[0108] As shown in FIGS. 8a-b, increasing the cell speed through
the constriction 15 can increase the delivery percentage and
delivery efficiency of the delivery material 30. It was found that
delivery efficiency varies linearly with cell speed, and that there
was a dosage dependent response.
[0109] As shown in FIG. 9, the incubation time of a cell in the
solution buffer 25 after the cell passes to the constriction 15 can
have an effect on the overall delivery percentage of the delivery
material 30 to the cell 20. It was noted, however, that after a
certain amount of incubation time (approximately 2-3 minutes), the
delivery percentage was substantially unchanged. Based upon this
data, it is believed that the perturbations caused in the cell 20
after it passes through the constriction 15 are corrected within on
the order of about five minutes after the cell 20 passes to the
constriction 15. Additionally, and for reference, -1 minute
corresponds to the control group.
[0110] As shown in FIG. 10, it was observed that passing the cells
20 through the constriction 15 multiple times can have an effect on
the overall delivery percentage, but that it negatively affected
the overall viability of the cells 20. To generate this data, cells
were passed through the constriction 15, collected, and passed
through the device again within approximately 1 minute.
[0111] It has been observed that during the time the cells 20 are
perturbed (e.g., after passing through the constriction 15) that
material from within the cell can be extracted through the
perturbations. Thus, it has been found that when the cells 20 are
perturbed, that material can flow in and out of the cell 20. This
property means that the system 5 can be used as a method of
sampling intracellular material without lysing the cell. The
perturbations in the cell membrane will preferably result in an
outflow of macromolecules from the cytoplasm and, thus, can be used
to probe the composition of the cytoplasm.
[0112] As shown in FIG. 11, stable green florescent protein (GFP)
expressing HeLa cells were treated in the presence of GFP silencing
siRNA (Ambion, U.S.A) and analyzed by FACS (FACS Canto II, BD
Biosciences, U.S.A.) at 48 hours for fluorescence knockdown. The
results in FIG. 11 indicate a >40% knockdown of gene
expression--a result comparable to that of commercial reagents such
as Lipofectamine 2000 (Invitrogen, U.S.A). Scrambled siRNA
controls, also in FIG. 11, indicate that this knockdown is not
caused by the deformation process itself.
[0113] As shown in FIGS. 13-14 and, the squeeze dimensions can have
an effect on the overall delivery efficiency of the delivery
material 30. For example, FIGS. 13-14 show that as the operating
pressure is varied (e.g., by varying the length and/or width of the
constriction 15) the overall delivery efficiency varies somewhat
(FIG. 14 relates to the delivery of quantum dots (nanopartices)
under different conditions). Furthermore, as shown in FIGS.
18a-18b, the estimated cell speeds can have an effect on the
overall viability and delivery efficiency of the delivery material
30. For example, FIG. 18a shows that as the operating speed is
varied, the overall delivery efficiency varies somewhat.
Additionally, FIG. 18b shows that as operating speed is varied, the
viability of the cells can vary somewhat. These figures show that a
change in constriction length can enhance delivery while minimally
impacting viability. Additionally, larger molecules enter the cell
at a lower rate after constriction than smaller molecules. This
intracellular delivery method described herein is "universal" in
that it works for many different type of materials and cells.
Further, the membrane disruptions induced by this device can be
typically at least .about.100 nm in size, although other size
disruptions are possible.
[0114] Referring to FIG. 12, in one implementation, the
concentration gradient between the solution buffer 25 and the
cytosol can be controlled to predictably control the amount of
delivered material. Localized delivery methods that expose the
cells 20 to a concentrated cloud of macromolecules after the cells
20 have been porated by the constriction can be used. Any such
localized delivery method, however, should account for the
estimated perturbation resealing time to ensure proper function.
This can be implemented by incorporating a "micronozzle"
perpendicular to the channel that delivers a high concentration of
the payload to the vicinity of the cell membrane (illustrated in
FIG. 6A). Preferably, the micronozzle can be located at and/or near
the constriction 15. Such an approach could allow supplementation
of the diffusive delivery mechanism with a convective component
thus enabling more precise cell loading with higher concentrations.
Preferably, the injection takes place while the cell 20 is in a
high concentration area of the constriction 15. A localized
technique has the added advantage of conserving valuable delivery
materials because it is then not necessary to maintain a high
concentration throughout the buffer.
[0115] Referring to FIG. 16a, a series of micropillars 100 can be
used to apply pressure to the cells 20 such that a perturbation is
caused. In this implementation, the cells 20 are forced through a
constricting pillar array in such a manner that pressure is applied
to the cells 20.
[0116] Referring to FIG. 16b, compression plates 105 can be used to
apply pressure to the cells 20 such that a perturbation is caused.
In this implementation, the compression plates 105 can be
controlled such that pressure is applied to the cells 20 for a
predefined amount time. The compression plates 105 can be
configured such that one or both plates move to apply pressure to
the cells 20. An additional sets of compression plates 105 can also
be supplied such that the cells 20 are substantially
surrounded.
[0117] Referring to FIG. 16c, buffer additives 115 (or bulking
materials bound to the cell surface) can be used to simulate
squeezing as the cell 20 passes through a constriction 15 that is
larger than the diameter of the cell 20. For example, simulated
constriction due to interference by the buffer additives 115 is
possible. Examples of buffer additives 115 include micro or
nanoparticles (e.g., polymer based, lipid based, ceramic based,
metallic, etc.). These particles are labeled with a cell binding
ligand such as an antibody, DNA sequence, peptide or small
molecule, although this is not required.
[0118] Referring to FIG. 16d, beads 120 can be used to compress the
cell 20. For example, magnetic and/or electrostatic force can be
used to apply pressure to the cell 20, or in the case of FIG. 16e,
to pull the cell 20. Preferably, the force applied to the cell 20
is sufficient to cause a perturbation.
[0119] Referring to FIG. 16f, multiple fluid streams 125 can be
directed in such a manner that compression (or rapid transitory
shearing) of the cell 20 is caused. For example, the multiple fluid
streams 125 can be fired in such a manner as that they approach or
impinge upon one another. As the cells 20 pass through the multiple
fluid streams 125, force can be applied to the cells 20 such that a
perturbation in the membrane of the cell 20 is caused.
Alternatively, cells can be fired through a narrow slit-like nozzle
to facilitate delivery.
[0120] The system 5 can be a standalone system, such as that shown
in FIG. 7, although other configurations are possible. For example,
the system 5 can be implanted in vivo in a patient for local
intracellular delivery, and or be incorporated ex vivo in a machine
for treatment of cells before returning the cells to the
patient.
[0121] In addition to its delivery advantages described herein, the
microfluidic nature of the system enables one to exercise precise
control over delivery conditions, pretreatment and subsequent
characterization of cells. For example, the system may be
implemented in series with a Fluorescence Activated Cell Sorting
(FACS) module. This can enable the delivery and sorting of the
desired cells on the same system, in real-time. Various
pretreatment and post-sort assaying techniques can also be
deployed, thus enabling the development of continuous,
high-throughput assays for drug screening and diagnostics.
[0122] The delivery efficiency of a payload delivered to target
cells is determined by subjecting a control population of target
cells to a payload as well as a population having undergone
treatment by a microfluidic device. The control sample is exposed
to the same delivery solution, at the same concentration, for at
least the same amount of time as the cells treated by the device.
To compensate for surface binding, endocytosis, and other effects
such as autoflourescence, a delivered region is defined such that
only the top 1-5% of live control cells fall into this region. The
delivery efficiency of a sample thus corresponds to the percentage
of live cells that are in the delivered region. For example, FIG.
43A is an intensity histogram from flow cytometry of a control
population that is exposed to cascade blue conjugated 3 kDa
dextran. FIG. 43B is an intensity histogram from flow cytometry of
cells that have been subjected to a 30 .mu.m-6 .mu.m device. The
defined delivered region is the unshaded region in both 43A and
43B.
[0123] In operation, referring to FIG. 17, with further reference
to FIGS. 1-3, a process 1000 for performing intracellular delivery
the system 5 includes the stages shown. The process 1000, however,
is exemplary only and not limiting. The process 1000 may be
altered, e.g., by having stages added, removed, altered, or
rearranged.
[0124] At stage 1005, the cells 20 are suspended solution buffer 25
along with delivery materials 30. Typical cell concentrations can
range from 10.sup.4 to 10.sup.9 cells/ml. Delivery material
concentrations can range from 10 mg/ml to 0.1 ug/ml. The delivery
material may be added to the cell buffer before or immediately
after delivery depending on the desired setup given that the
injuries/pores remain open for 1-5 minutes. The solution buffers
may be composed of a number of salts, sugars, growth factors,
animal derived products or any other component necessary for proper
cell proliferation, maintaining cell health or induction of cell
signaling pathways. Additional materials may also be added to the
solution buffer 25. For example, surfactants (e.g., pluronics)
and/or bulking materials can be added to the solution buffer
25.
[0125] At stage 1010, the solution buffer 25 including the cells 20
and the delivery materials 30 are passed through the channel 10 of
the system 5. The solution buffer 25 can pass through the channel
10 using gravity, or can be assisted by other methods. For example,
pressure can be applied to the solution buffer 25 on the entrance
side of the channel 10 (e.g., using a gas cylinder and/or
compressor), and/or a vacuum can be applied by a vacuum pump on the
exit side. Additionally, displacement based flow systems can also
be used.
[0126] As the individual cells 20 pass through the constriction 15,
a pressure is momentarily applied to the cell 20 by the solid
construction of the constriction 15 causing perturbations such as
holes to develop in the cell membrane such that the delivery
materials 30 can be delivered to the inside of the cell 20. The
amount and/or duration of the pressure applied to the cell 20 can
be varied by adjusting the dimensions of the constriction 15, the
velocity at which the cell 20 passes through the constriction 15,
and/or by adjusting the shape of the constriction 15. In one
configuration, approximately 5,000-20,000 cells/second pass through
the constriction 15, and each cell is constricted for approximately
100 .mu.s.
[0127] The system 5 can include one or more of the channels 10. For
example, the system 5 can include 50-100 of the channels 10 that
are arranged in a parallel configuration. Using a parallel
configuration can reduce the consequences of a clog developing in
one or more of the channels 10, and can increase the overall
throughput of the system 5. Additionally, the system 5 can include
one or more of the channels in series with one another.
[0128] At stage 1015, after the cells 20 pass through the
constriction 15, the cells are allowed to incubate/recover by
sitting in the solution buffer 25. During this time, the cells 20
will intake some of the delivery materials 30 is present in the
solution buffer 25 through the perturbations in the cell membrane.
One mechanism of intake is diffusion-based, because larger
molecules appear to be absorbed at a slower rate than smaller
molecules. Preferably, the cells 20 are allowed to incubate/recover
in the solution buffer 25 for on the order of 2-5 minutes, although
other durations are possible. During the time that the cells 20 are
incubating/recovering in the solution buffer 25, material from
inside the cell 20 may also release from the cell into the solution
buffer 25. During the incubation/recovery period, certain
conditions can be controlled to ensure that delivery quantities of
the delivery materials 30 are consistent across the cell
population. For example, post-constriction, convective delivery
mechanisms that impinge delivery material onto the
incubating/recovering cell can be used.
[0129] Optionally, at stage 1020, after the cells have
incubated/recovered, the cells can be washed to remove the solution
buffer. Preferably, the washing occurs after the time period
required for the perturbations to be repaired, although the washing
can occur at other times in order to control the amount of delivery
materials 30 absorbed by the cells.
Example 1--Delivery of Functional Engineered Nanoparticles
[0130] Engineered nanomaterials have immense potential as live cell
imaging tools, therapeutic molecular delivery agents, or even as
ways to manipulate live cells with external handles such as light
or magnetic fields. (Howarth, M., et al. Monovalent, reduced-size
quantum dots for imaging receptors on living cells. Nature Methods
5, 397-399 (2008)). However, much of these potential applications
require that nanomaterials be delivered into the cell cytosol. Most
nanoparticles, such as QDs, need to be passivated with a polymer
that renders the nanoparticles soluble in aqueous media, and this
generally prevents them from passively diffusing across the cell
membrane. Microinjection of nanoparticles is considered impractical
due to specialized equipment requirement and low throughput while
electroporation causes QD aggregation inside the cell. Therefore,
most attempts to deliver QDs into the cell cytoplasm have relied on
QDs being endocytosed by the cell and escaping from the endosome.
Prior to the current subject matter, it was not possible to deliver
QDs into cell cytoplasm in a satisfactory and scalable manner. They
system provides a solution to this delivery problem of earlier
approaches.
[0131] The microfluidic device is combined with a new generation of
recently described biologically compatible QDs. (Liu, W., et al.
Compact biocompatible quantum dots via RAFT-mediated synthesis of
imidazole-based random copolymer ligand. JACS 132, 472-483 (2010)).
The QDs used throughout example 1 were coated with a poly-imidazole
ligand comprised of multiple metal-chelating imidazole groups and
multiple water-solubilizing, passivating poly(ethylene) glycol
(PEG).
[0132] For cytosolic delivery of QDs, cells were aliquoted into a
PBS solution containing QD. The cell-QD solution was pipetted into
the microfluidic device and the solution was driven through the
channels at constant pressure, followed by a 5 min incubation
period. After this incubation period, excess QDs were separated by
centrifugation. For the control population, the cell-QD solution
was placed in the microfluidic device and the cells were exposed to
the QD solution for an amount of time equivalent to the cytosolic
delivery protocol.
[0133] FIG. 19 is an overlay of transmission and confocal
fluorescence images, followed by z-section confocal fluorescence
images of treated cells delivered with QDs using the current
subject matter. FIG. 19 illustrates (top) immediately after
treatment (i.e. delivery) and (bottom) after 48 h incubation at
37.degree. C. and 5% CO.sub.2. The diffuse staining pattern is
constrained to the cytoplasm and the nanoparticles appear not to
enter the nucleus (dark region within the cell). Scale bar is 10
.mu.m. The particular free poly-imidazole ligand that coated the
QDs imaged in FIG. 19 had no functionality other than providing
biocompatibility through PEG groups. Confocal microscopy images
show that HeLa cells, detached and round after flowing through the
microfluidic device, have diffuse cytoplasmic QD staining
throughout different z-sections of the cell (FIG. 19, top). The
diffuse staining persists even after 48 hours, following incubation
and adherence of the cells at 37.degree. C. in 5% CO.sub.2 (FIG.
19, bottom). The diffuse QD fluorescence is dimmer at 48 hrs,
likely due to cell division (FIG. 19). The device delivered QDs
(.about.13 nm hydrodynamic diameter) into .about.40% of the live
cell population at a throughput rate of .about.10,000 cells/s. FIG.
20A illustrates delivery efficiency into HeLa cell cytosol upon
current subject matter treatment with QDs coated with PIL. Cell
viability was >80% as measured by flow cytometry. FIG. 20B
illustrates viability of HeLa cells upon delivery of plain QD535 by
the current subject matter, as measured by propidium iodide
staining and flow cytometry measurement. The viability of treated
cells as measured by flow cytometry, the diffuse staining on the
confocal images, and the cell's ability to adhere are consistent
with delivery of QDs into the cytoplasm of a live cell.
[0134] To confirm that the fluorescence indeed arises from QDs
delivered to the cytosol as opposed to QDs sequestered in
endosomes, the nanoparticle was designed to change its emission
profile upon interaction with the reducing environment of the
cytosol. The reduction potential inside the cell cytoplasm is -260
to -220 mV and is primarily dictated by the maintenance of high
concentrations (5-10 mM) of the tripeptide glutathione. Therefore,
by measuring the fluorescence of a QD-dye construct whose emission
changes when exposed to the cytosolic environment, the localization
and chemical accessibility of the delivered nanoparticles can be
determined. A QD-dye was constructed comprising of a green emitting
QD (.lamda.emission=541 nm) that acts as an energy donor to a
carboxy-X-Rhodamine (Rox) dye (.lamda.emission=610 nm), conjugated
through a reducible disulfide bond.
[0135] FIG. 21 illustrates construct design, absorbance, and
stability in various media. At 2100 is a schematic of the PIL prior
to conjugation with the dye and coating the QDs (left), and of the
resulting QD-disulfide-Rox construct (right) (image not to scale).
At 2110 is the absorbance spectrum of the QD-disulfide-dye
construct. Excitation at 488 nm and at 405 nm provided exclusive
absorption by the QDs throughout the experiment. At 2120 is the
stability of fluorescence energy transfer from QD to Rox for the
construct in full culture media at 37.degree. C. and 5% CO.sub.2,
demonstrating that the disulfide bond is not cleaved in the
extracellular environment. Plot 2130 illustrates cleavage of the
disulfide bond by the cytosolic reductant glutathione, as shown by
the recovery of QD fluorescence. At 2140, recovery of QD
fluorescence upon treatment by the non-thiol reductant
tris(2-carboxyethyl) phosphine is shown, further supporting the
cleavage of the disulfide bond.
[0136] Thiol groups that were incorporated into the PIL formed
disulfide bonds with thiolated Rox dyes. The absorbance spectrum of
the purified construct has absorbance features of both QD and Rox
(2120) at an average of 13 Rox dyes per QD, effectively quenching
the QD fluorescence (2130). This construct serves as an
irreversible sensor of the specific reducing environment in the
cytosol. When the QD is selectively excited by a laser at 488 nm
(microscopy) or 405 nm (flow cytometry) while the disulfide bridges
are intact, the construct undergoes fluorescence resonance energy
transfer (FRET) so that Rox emission in the red dominates. In a
solution assay, the cellular reductant glutathione cleaves the
disulfide bridges, releasing Rox dyes and allowing the QD
fluorescence to recover (2140). The non-thiol based reductant
tris-(2carboxyethyl) phosphine also allows QD fluorescence
recovery, indicating that the release of Rox from the QD surface is
not via PIL displacement by glutathione (2140). Rox fluorescence
may not completely disappear due to some of the disulfide bridges
being sterically hindered by long PEG groups on the PIL, and due to
some small amount of non-specific interaction between the dye and
the QD surface.
[0137] Changes in the fluorescence profile of the construct, as
measured by flow cytometry and confocal microscopy, confirm the
delivery of QD-disulfide-Rox constructs to the cell cytoplasm. When
exposed to the reducing cytosolic environment, the cleavage of the
disulfide bonds disrupts the FRET process from the QD to the dye.
Therefore, upon exclusive excitation of the QD, QD channel
fluorescence increases while Rox channel fluorescence decreases
with time. Live HeLa cells were treated by the microfluidic device
in a solution with a high concentration of QD-disulfide-Rox,
incubated for 5 minutes, and washed to remove excess QDs before
adding cell culture media (i.e. the treated cells). Control cells
were incubated with QD-disulfide-Rox for 5 minutes instead of being
treated by the microfluidic device, and washed before being placed
in cell culture media. The Rox and QD channel fluorescence of these
treated and control cells were observed by both confocal microscopy
and flow cytometry.
[0138] FIGS. 22A and 22B illustrate live cell confocal microscopy
images and fluorescence intensity analysis demonstrating
cytoplasmic staining and chemical accessibility of QD surface. FIG.
22A illustrates images of treated cells (top) and control cells
(bottom). The appearance of diffuse green fluorescence is present
only in treated cells. Scale bar is 10 .mu.m. FIG. 22B illustrates
a change in intensity as a function of time in the green and red
channels. Because n<20 at each time point, fluctuations in total
average fluorescence were corrected by normalizing to the 0 h time
point.
[0139] Under the confocal microscope, the diffuse fluorescence that
appears across the cytoplasm of treated cells progresses from
strongly red to strongly green (as shown in FIG. 22A). Control cell
images show some non-specific binding on the outer membrane as
demonstrated by the ring-shaped fluorescence, and there is no
increase in green channel signals. These effects are consistent
with the expected cleavage of cytosolic disulfide bonds which
reduce the FRET effect. In FIG. 22B, the line graph plots the
average QD and Rox channel intensity per cell after correcting for
cell-to-cell differences in delivered fluorescent material by
normalizing for total fluorescence, for treated and control cells
and autofluorescence. For treated cells, the graph shows a cross
over between 2-4 hours of incubation where the QD fluorescence
rises above the Rox fluorescence. Interestingly, the treated cell
Rox signal is shown to stabilize above autofluorescence levels
after 9 hours. This is consistent with results from solution
assays, where some FRET remained after reduction. The observed
diffuse staining and increase in QD signal and reduction in Rox
signal strongly support cytosolic delivery and subsequent disulfide
bond cleavage. The QD fluorescence in control cells, quenched by
FRET to the Rox, appears indistinguishable from autofluorescence.
The control cells display some Rox fluorescence above
autofluorescence at early time points, which then steadily
decreases. This can be attributed to non-specific interactions
between QD-S--S-Rox and the surface of the cell, followed by
re-solvation of the constructs into the medium.
[0140] FIGS. 23A-23D illustrate flow cytometry measurements of
average cell fluorescence and viability. At 2300 is average
fluorescence of QD (left) and Rox (right) per cell, showing an
increase in QD fluorescence only in treated cells. Rox fluorescence
in both treated and control cells is at autofluorescence levels by
the 24 h time point. At 2310 is a histogram of the distribution of
fluorescence intensities among treated and control cells at select
time points, in the QD channel (left) and Rox channel (right). QD
delivery is estimated to have occurred in at least 35% of the cell
population. Grey areas are meant to guide the eye in the movement
of fluorescence intensity histogram peaks. At 2320 illustrates
viability of control and treated cells as measured by propidium
iodide.
[0141] The flow cytometry measurements illustrated in FIGS. 23A-23D
confirm that the QD-disulfide-Rox constructs can interact with the
cytosolic environment. Flow cytometry measurements were recorded on
all live cells, encompassing both delivered (.about.35% of the
treated cell population) and undelivered cells. At 2100 the average
fluorescence per cell of the treated and control populations is
illustrated. The average QD fluorescence rises initially for the
treated cells, peaking at .about.9 hrs and falling gradually
thereafter, in contrast to the QD fluorescence of the control cell
population, which stays comparable to autofluorescence levels. This
is consistent with the cytosolic reduction of disulfide bridges
between the QD and dye inside the treated cells followed by
dilution of fluorescence constructs by cell division. The Rox
fluorescence for both the treated and control cells start high and
drop within the first 2 hrs. This drop is attributed to the
re-solvation into the medium of particles that had become bound to
the cell surface during incubation. The average Rox fluorescence in
the treated cell population appears similar to control cells due to
the presence of undelivered cells within the treated population.
The presence of both delivered and undelivered cells within the
treated population can be distinguished in the histograms of QD and
Rox intensity shown at 2310. With increasing time, the fluorescence
histograms become bimodal for treated cells but stay unimodal for
control cells. QD fluorescence rises with time in a subset of the
treated cell population (2100), further supporting the disruption
of the FRET process in the cytosol of treated and delivered cells.
Rox fluorescence decreases overall as membrane-bound constructs are
re-solvated into the medium, but a subset of the treated cell
population retains Rox fluorescence. This is consistent with the
incomplete reduction of QD-S--S-Rox bonds observed in confocal
microscopy. The viability of the treated cell population, as
measured by propidium iodide staining, is within 10% of the control
population at all time points (2130). The cell viability of >90%
relative to the control group compares favorably to alternative
methods such as electroporation and polymer-based methods, which
have yielded post-treatment viabilities as low as 40-60%.
[0142] FIG. 24 illustrates epifluorescence imaging of unaggregated
single QDs within the cell cytosol after device treatment with a 10
nM QD solution (top), and blinking traces of the three QDs labeled
2400, 2410, and 2420 with autofluorescence. QD blinking traces
appear to be non-binary due to long acquisition bin times (500 ms).
Scale bars are 10 .mu.m.
[0143] The QD delivery platform also enabled single molecule
imaging by delivering unaggregated QD-disulfide-Rox constructs, as
the observed emission intermittency is consistent with single QDs.
For this experiment, QD-disulfide-Rox constructs were delivered
into the cytosol followed by a 10 hour incubation and imaged on an
epifluorescence microscope. The 10 hour incubation ensured that the
QD fluorescence from inside the cytosol has recovered via disulfide
bond reduction; epifluorescence microscopy was used to ensure that
enough photons are collected. Several blinking QDs were observed
when cells were treated by the current subject matter at low QD
concentrations (FIG. 24). Intensity traces of blinking QDs in the
cytosol, shown at 2400, 2410, and 2420, appear non-binary as a
result of long acquisition bin times (500 ms). Translational cell
movements were deemed minimal during the time frame of the
acquisition (.about.1 min). These data demonstrate the capability
of observing single molecule events within the cell cytosol by
delivering QDs as fluorescent labels using the current subject
matter.
[0144] Example 1 demonstrates nanoparticle delivery into cell
cytosol according to an embodiment of the current subject matter.
By observing the cleavage of QD-disulfide-Rox by cytosolic
reductants, it has been shown that the nanoparticle surface
interacts with cytosolic components. Embodiments of the current
subject matter enables delivery of QDs into cell cytoplasm at high
throughput without any cell penetrating or endosome escaping
ligands, while conserving cell viability and QD integrity. The
delivery efficiency of 35% may be further increased by increasing
the number of microfluidic constrictions, changing constriction
dimensions, or increasing the number of treatment cycles. Unlike
most of the current cell penetrating peptide or positive
charge-assisted delivery methods, the current subject matter does
not require dual conjugation of an intracellular delivery handle
and a cytosolic protein-targeting handle on the same nanoparticle.
By dispensing the need for the former, mitigation of the concerns
of cross-reactivity, unequal reactivity efficiencies of conjugation
strategies, and conjugation stoichiometry can be achieved.
Therefore, significant flexibility in QD construct design is
garnered, paving the way for intracellular protein labeling and
tracking. The methods are useful for the delivery of many
fluorescent nanomaterials with complex designs that target
intracellular proteins and organelles through proven
protein-targeting strategies such as, but not limited to,
streptavidin-biotin, HaloTag-chloroalkane, and sortase tagging.
[0145] In example 1, all chemicals were obtained from Sigma Aldrich
and used as received unless indicated otherwise. Air sensitive
materials were handled in an Omni-Lab VAC glovebox under dry
nitrogen atmosphere with oxygen levels <0.2 ppm. All solvents
were Spectroscopic or reagent. Aromatic ring-bearing compounds were
visualized on TLC using a hand-held UV lamp and KMnO4.
Amine-bearing compounds were visualized on TLC using a Ninhydrin
stain. Flash column chromatography was performed on a Teledyne Isco
Combi Flash Companion. HeLa cells were purchased from ATCC and all
cell medium materials were purchased from Mediatech unless
indicated otherwise.
[0146] In example 1, 1H NMR spectra were recorded on a Bruker DRX
401 NMR Spectrometer. MS-ESI was performed on a Bruker Daltonics
APEXIV 4.7 FT-ICR-MS machine. UV-Vis absorbance spectra were taken
using an HP 8453 diode array spectrophotometer. Photoluminescence
and absorbance spectra were recorded with a BioTek Synergy 4
Microplate Reader. Polymer molecular weights were determined in DMF
solution on an Agilent 1100 series HPLC/GPC system with three PLgel
columes (103, 104, 105 .ANG.) in series against narrow polystyrene
standards. Dye derivatives were purified using Varian ProStar Prep
HPLC system. Modified polymer was purified using GE Healthcare's
PD-10 columns packed with Sephadex.TM. G-25M. Ligand exchanged QDs
were purified by centrifugation dialysis with Millipore Amicon
Ultra 30K cut-off centrifugal filters and by GFC on AKTAprime Plus
chromatography system (Amersham Biosciences) equipped with a
self-packed Superdex 200 10/100 glass column. Flow cytometry
measurements were made on LSR Fortessa (BD Biosciences).
[0147] In example 1, CdSe cores with 478 nm first absorption peak
were synthesized using a previously reported method (1). To
summarize, 0.4 mmol (54.1 mg) of CdO, 0.8 mmol (0.2232 g) of TDPA,
9.6 mmol (3.72 g) of TOPO were placed in 25 mL round bottom flask.
The solution was degassed for 1 hr at 160.degree. C. and heated to
300.degree. C. under argon until the CdO dissolved and formed a
clear homogenous solution. This was followed by putting the
solution under vacuum at 160.degree. C. to remove evolved water.
The solution was reheated to 360.degree. C. under argon and a
TOP-Se solution (1.5 mL of 1.5M TOP-Se in 1.5 mL of TOP) was
rapidly added to give CdSe cores with the first absorption feature
at 478 nm.
[0148] CdS shells were deposited on CdSe cores via modification of
previously reported procedures (2). Cores isolated by repeated
precipitations from hexane with acetone were brought to 180.degree.
C. in a solvent mixture of oleylamine (3 mL) and octadecene (6 mL).
Cd and S precursor solutions were then introduced continuously at a
rate of 4 mL/hr. The Cd precursor consisted of 0.33 mmol Cd-oleate
and 0.66 mmol oleylamine in a solvent mixture of octadecene (1.5
mL) and TOP (3 mL). The S precursor consisted of 0.3 mmol
hexamethyldisilathiane [(TMS)2S] in 6 mL TOP. Addition of a total
of 3 monolayers each of Cd and S yielded QDs with emission at 541
nm and a quantum yield of 60% when diluted in octane. The
extinction coefficient of CdSe(CdS) was calculated using the
extinction coefficient of CdSe cores from the literature (3) and
assuming that 95% of the CdSe cores were retained during the
overcoating step.
[0149] In example 1, the silicon chip was fabricated using
photolithography and deep reactive ion etching techniques. The
resulting etched silicon wafer was cleaned (with H2O2 and H2SO4) to
remove debris, oxidized to produce a glass surface, and bonded to a
Pyrex wafer before being diced into individually packaged devices.
Each device was then individually inspected for defects prior to
use.
Example 2--Delivery of Macromolecules
[0150] Intracellular delivery of macromolecules is a critical step
in therapeutic and research applications. Nanoparticle mediated
delivery of DNA and RNA, for example, is useful for gene therapy,
while protein delivery is being used to affect cellular function in
both clinical and laboratory settings. Other materials, such as
small molecules, quantum dots, or gold nanoparticles, are delivered
to the cytosol for purposes ranging from cancer therapies to
intracellular labeling and single molecule tracking.
[0151] To demonstrate the versatility of the technique, model
dextran molecules were delivered to several cell types: DC2.4
dendritic cells, newborn human foreskin fibroblasts (NuFF) and
mouse embryonic stem cells (mESC) attained delivery efficiencies of
up to 55%, 65% and 30% respectively. Initial experiments also
showed successful delivery in primary lymphocytes, macrophages and
dendritic cells derived from mice. Moreover, the technique did not
cause excessive cytotoxicity or induce stem cell differentiation.
Indeed all cell types were over 60% viable even at the highest
tested speeds. Device design and operating conditions had not been
previously optimized for any of the aforementioned cell types.
[0152] FIG. 25 illustrates experimental results showing that
delivery performance depends on cell speed and constriction design.
Constriction dimensions are denoted by numbers (e.g. 10 .mu.m-6
.mu.m.times.5) such that the first number corresponds to
constriction length, the second to constriction width and the third
(if present) to the number of constrictions in series per channel.
At 2500, delivery efficiency is shown and at 2510 cell viability 18
hours post treatment (measured by flow cytometry) is shown as a
function of cell speed for 40 .mu.m-6 .mu.m (.smallcircle.), 20
.mu.m-6 .mu.m (.quadrature.) and 10 .mu.m-6 .mu.m.times.5 (.DELTA.)
device designs. At 2520, delivery efficiency and 2530 cell
viability (measured by flow cytometry) is shown as a function of
speed in primary human fibroblasts (.quadrature.), DC2.4 dendritic
cells (.smallcircle.), and mouse embryonic stem cells (mESC)
(.DELTA.) treated by a 30 .mu.m-6 .mu.m device. Human fibroblast
and dendritic cells were analyzed 18 hours post-delivery. MESCs
were analyzed 1 hour post-delivery. All data points were run in
triplicate and error bars represent two standard deviations.
[0153] FIG. 26 illustrates scans 2600 of different horizontal
planes of a HeLa cell after the delivery of pacific blue conjugated
3 kDa dextran, as measured by confocal microscopy. Scans read from
top to bottom, then left to right where the top left is at z=6.98
.mu.m and bottom right is at z=-6.7 .mu.m. Scale bar represents 6
.mu.m. At 2610, live cell delivery efficiency of 10 .mu.m-6 .mu.m
(.quadrature.), 20 .mu.m-6 .mu.m (.smallcircle.), 30 .mu.m-6 .mu.m
(.DELTA.), and 40 .mu.m-6 .mu.m (.diamond.) devices is shown. The
time axis indicates the amount of time elapsed from initial
treatment of cells before they were exposed to the target delivery
solution. All results were measured by flow cytometry 18 hours
post-treatment. At 2620, average intensity of the delivered cell
population normalized by untreated cells to control for
autofluorescence. Fluorescein conjugated 70 kDa dextran (horizontal
lines) and pacific blue conjugated 3 kDa dextran (diagonal lines)
are delivered to the cell (cycles 1 and 3) and removed from the
cell (cycle 2) in consecutive treatment cycles. The control
represents cells that were only exposed to the delivery solution
and not treated by the device. All data points were run in
triplicate and error bars represent two standard deviations.
[0154] As previous nanoparticle and cell penetrating peptide
(CPP)-based delivery techniques exploit endocytotic pathways,
evidence is presented that rules out the influence of endocytosis
in the current subject matter delivery mechanism. FIG. 26
illustrates at 2600 confocal microscopy of cells treated with
pacific blue conjugated 3 kDa dextran demonstrate diffuse cytosolic
staining as opposed to the punctate characteristic one would expect
of endocytotic methods. Moreover, when delivery experiments are
conducted at 4.degree. C., a temperature at which endocytosis is
minimized, delivery efficiency is minimally affected by temperature
for both test payload materials, 3 kDa and 70 kDa dextran. This
data indicate that endocytosis is unlikely to be responsible for
delivery in this system.
[0155] Delivery kinetics over time were characterized. Cells were
treated by the current subject matter in the absence of delivery
material and subsequently exposed to pacific blue labeled 3 kDa
dextran at defined time intervals post-treatment. In this approach,
as the cells pass through the constriction their membrane is
disrupted; however, no measurable delivery occurs until they are
exposed to the labeled dextran. Thus, the delivery efficiency at
each time point would reflect the proportion of cells that remained
porous for that amount of time post-treatment. This method captures
the kinetics of pore formation/closure. The results indicate that
almost 90% of delivery occurs within the first minute after
treatment regardless of device design (2610). The observed
time-scale supports the pore formation hypothesis as previous works
on membrane repair kinetics have reported membrane sealing
occurring at about 30 s after an injury is induced. In contrast,
the recommended time-scale for endocytotic methods such as
nanoparticle and CPP mediated delivery mechanisms is on the order
of hours.
[0156] Since delivery of material through the membrane pores is
diffusive, material could be exchanged into and out of the cell
throughout the lifetime of the pore. Endocytotic or convective
mechanisms, on the other hand, must be unidirectional, i.e. only
facilitate transport of material into the cell. To demonstrate
bidirectional transport of material across the cell membrane an
experiment was conducted consisting of 3 delivery cycles. In the
first cycle, cells were treated in the presence of 3 kDa and 70 kDa
dextran, incubated for 5 min in the dextran solution and washed
twice with PBS. One third of the sample was retained and plated for
follow-up. In the second cycle, the remaining washed cells were
treated by the device again but in the absence of any delivery
material and incubated for another 5 min. Half of this sample was
plated for follow-up. In the third cycle, the remaining cells from
the second cycle were run through the device under the same
conditions as the first cycle (i.e. in the presence of dextran),
incubated for 5 minutes and washed twice in PBS. The cells were
analyzed by flow cytometry 18 hours after the experiment. The
changes in normalized fluorescence intensity demonstrate a net
diffusion of dextran into the cells during the first cycle, out of
the cells during the second, and back in during the third (2620).
These results are thus consistent with the diffusive delivery
mechanism.
[0157] FIG. 27 illustrates a simplified, 2D diffusion model
developed in a software package known as COMSOL Multiphysics that
simulates passive diffusion of material into a cell across a
porated membrane. COMSOL Multiphysics is a finite element analysis,
solver and Simulation software/FEA Software package developed by
COMSOL for various physics and engineering applications, especially
coupled phenomena, or multiphysics. At 2700, delivery/loss of
material is shown as a function of membrane diffusivity. Simulation
results indicating the percentage of material delivered/lost from
the cell as a function of membrane diffusivity when the material of
interest is in the buffer (.quadrature.) or in the cell
(.smallcircle.) at the time of poration. At 2710 is a graphical
representation of the simulated system and the concentration
gradient that forms across the membrane if material is delivered
from the buffer to the cell.
[0158] Using literature values for particle diffusivities inside
and outside the cell cytoplasm, the experimental results of FIG. 26
were qualitatively recreated with diffusion as the only mode of
mass transfer. Moreover, by fitting the experimental data to this
model, this technique delivers 10-40% of the delivery material in
the buffer into the cell cytosol. By comparison, CPP methods for
protein delivery are estimated to deliver only 0.1% of the buffer
material to the cytosol.
[0159] A particle's size (or hydrodynamic radius) affects its
diffusivity and its ability to enter membrane pores of a particular
size. Thus, this parameter affects delivery efficiency in the pore
formation/diffusion mechanism. In a series of experiments, test
payloads of 3 kDa, 10 kDa, 70 kDa, 500 kDa, and 2 MDa dextrans
conjugated to fluorescein or pacific blue were delivered.
Fluorescein labeled plasmids estimated at 3.1 MDa were also
delivered. These model molecules were selected based on their
similarity in molecular weight to delivery materials of interest. 3
kDa-10 kDa dextran, for example, are of similar size to some short
peptides or siRNA, while the 70 kDa-2 MDa range mimics the size of
most proteins and some small nanoparticles.
[0160] FIG. 28 illustrates results of a two-tiered delivery of
material. At 2800, live cell delivery efficiency, as a function of
speed, for HeLa cells treated with Pacific Blue conjugated 3 kDa
(.quadrature.), fluorescein conjugated 70 kDa (.smallcircle.) and 2
MDa (.DELTA.) dextran is shown. This experiment was conducted with
a 10 .mu.m-6 .mu.m.times.5 chip. All data points were run in
triplicate and error bars represent two standard deviations. 2810
and 2820, illustrate histogram overlays of flow cytometry data for
HeLa cells that are untreated (red), treated at 700 mm/s (green),
treated at 500 mm/s (orange), treated at 300 mm/s (light blue), or
only exposed to the delivery material (control, dark blue). The
delivery material consisted of pacific blue conjugated 3 kDa
dextran (2810) and fluorescein conjugated 70 kDa dextran
(2820).
[0161] The experiments have shown that molecules larger than 70 kDa
have a different delivery profile relative to 3 kDa dextran (2800).
The device produced a two-tiered delivery where a 10 .mu.m-6
.mu.m.times.5 device operated at 500 mm/s, for example, enables
over 90% of live cells to receive the 3 kDa molecules, while about
50% receive the larger 70 kDa and 2 MDa molecules.
[0162] The histograms corresponding to these flow cytometry data
indicate that 3 kDa dextran delivery produces two distinct peaks
(2810). In the first subpopulation, cells exhibit mild delivery
levels as observed by a peak shift relative to controls (controls
account for endocytosis and surface binding as described earlier
for the 0 mm/s data points) with a 2-6.times. increase in average
fluorescence intensity. In the second population, cells exhibit
enhanced delivery levels corresponding to a 20-100.times. increase
in average fluorescence intensity relative to controls. This effect
may indicate that the latter subpopulation of cells was more
severely porated than the former, hence enabling an almost
10.times. increase in material influx. Indeed, as illustrated by
the 300 mm/s, 500 mm/s and 700 mm/s curves, increasing the
treatment severity, by increasing operating speeds, appears to
increase the proportion of cells with enhanced delivery. One
observes a similar characteristic for the delivery of larger 70 kDa
dextran molecules (2820). The effect is less pronounced, however,
as the lower particle diffusivity and possible size exclusion
effects reduce the overall quantity delivered. The mild delivery
population (first peak) only shows a 1.5-2.times. increase in
average fluorescent intensity as compared to the 2-6.times.
observed in the 3 kDa case. This effect could account for the
discrepancy in the delivery data in 2800 as in the case of larger
molecules the mild delivery population could be difficult to
distinguish from controls based on the present definition of
delivery. As a result, for larger molecules, such as the 70 kDa and
2 MDa dextran, one largely measures the second, enhanced delivery
population
[0163] To verify that material delivered by rapid mechanical
deformation is active and biologically available in the cell
cytosol a series of experiments were conducted delivering test
payload GFP silencing siRNA (Ambion, USA) to HeLa cells expressing
destabilized GFP. Dose dependent and sequence specific GFP
knockdown (up to 80%) at 18 hours post-treatment was observed. The
gene knockdown response to cell speed and device design was
consistent with dextran delivery experiments such that higher
speeds and multiple constriction designs yielded greater gene
knockdown. Lipofectamine 2000 was used as a positive control in
these experiments. Device design and operating parameters were not
optimized for siRNA delivery prior to performing these
experiments.
[0164] FIG. 29 illustrates data relating to SiRNA, protein, and
nanoparticle delivery. At 2900, gene knockdown is illustrated as a
function of device type and cell speed, in destabilized GFP
expressing HeLa cells 18 hours after the delivery of anti-eGFP
siRNA using a 10 .mu.m-6 .mu.m.times.5 chip at a delivery
concentration of 5 .mu.M. Lipofectamine 2000 was used as a positive
control. At 2910, delivery efficiency of fluorescein labeled 70 kDa
and pacific blue labeled 3 kDa dextran by rapid mechanical
deformation and electroporation is shown. Each dextran type was at
a concentration of 0.1 mg/ml in the delivery solution. At 2920,
fluorescein labeled bovine serum albumin (.smallcircle.) delivery
efficiency, pacific blue conjugated 3 kDa dextran (.quadrature.)
delivery efficiency, and cell viability (.DELTA.), is shown as a
function of speed, using a 30 .mu.m-6 .mu.m device. At 2930,
fluorescent micrographs of HeLa cells immediately after delivery of
antibodies to tubulin with an Alexa Fluor 488 tag are shown. Scale
bars at 5 .mu.m. At 2940 and 2950 is shown tunneling electron
microscopy (TEM) images of gold nanoparticles (some indicated by
arrows) in cells fixed .about.1 s after treatment by a 10 .mu.m-6
.mu.m.times.5 device. Scale bars at 500 nm. All data points were
run in triplicate and error bars represent two standard
deviations.
[0165] In further experiments, cytosolic delivery potential is
explored for previously challenging applications, such as protein
and nanoparticle delivery. To compare the performance of the
current subject matter to commercially available methods, 3 kDa and
70 kDa dextran was delivered, as a protein model, to human
fibroblasts using the current subject matter and a Neon
electroporation system (Invitrogen). Results indicate that rapid
mechanical deformation provides a 7-fold increase or greater in
delivery efficiency for such macromolecules (2910). To translate
the method to protein delivery, a 30 .mu.m-6 .mu.m channel diameter
device was used to deliver fluorescein labeled bovine serum albumin
(BSA) to HeLa cells at up to 44% efficiency while maintaining
viabilities above 90% (2920). Alexa Fluor 488 labeled antibodies to
tubulin (BioLegend) were also delivered to HeLa cells after
treatment with a 30 .mu.m-6 .mu.m device using a 0.25 mg/ml
antibody concentration (2930). The diffuse staining indicates that
the material is not trapped in endosomes and hence is suitable for
live cell antibody staining of cellular structures in the cytosol.
Apolipoprotein E was also delivered successfully using this
technique.
[0166] For nanoparticle delivery, TEM images of cells fixed
.about.1 s after deformation (2940 and 2950) demonstrate the
delivery of PEG1000 coated, 15 nm gold nanoparticles. The gold
nanoparticles appear to be mostly un-aggregated and were not
visibly sequestered into endosomes. In these images, evidence for
various defects in the cell cytoplasm responsible for delivery were
observed. High throughput has been demonstrated, non-cytotoxic
delivery of quantum dots directly to the cell cytosol--a goal that
previous techniques have struggled to achieve. In these
experiments, quantum dots with a Rox dye bound to their surface
were delivered by rapid mechanical deformation and observed over
time. These data yielded a minimum estimated delivery efficiency of
35% and confirmed that the delivered quantum dots were in the
cytosol and chemically accessible to the intracellular
environment.
[0167] Transient pores are formed by rapid mechanical deformation
of a cell as it passes through a microfluidic constriction. Data
supports this mechanism by demonstrating diffuse cytosolic staining
(FIG. 26 at 2600), siRNA functionality (FIG. 29 at 2900) and the
bidirectional movement of material across the porated membrane
(FIG. 26 at 2920). A number of parameters have been identified,
such as constriction dimensions, number of constrictions in series
and cell speed that affect delivery efficiency and viability (FIG.
25). These parameters may thus be used to optimize device design
for individual applications based on cell type and the size of the
delivery material.
[0168] In example 2, this technique is based on the microfluidic
system, which can be incorporated into a larger integrated system
consisting of multiple pre-treatment steps prior to delivery and
analytical or sorting steps post-treatment. At an average
throughput rate of 10,000 cells/s, the delivery device can, for
example, be placed in-line with a flow cytometry machine to sort
cells or perform other analytical tasks immediately after
delivery.
[0169] The devices, systems, and methods described herein provide a
number of 30 potential advantages over existing methods. Similar to
electroporation and microinjection, it is a poration based
mechanism and hence does not rely on exogenous materials, chemical
modification of payloads or endocytotic pathways. In contrast to
electroporation, however, it does not rely on electrical fields
which have had limited success in protein delivery, can damage some
payload, or cause cytotoxicity. Indeed current results have
demonstrated relatively high viability in most applications and
sensitive payloads, such as quantum dots, appear to be undamaged.
The current subject matter thus provides significant advantages in
areas such as the labeling and tracking of cytosolic material where
quantum dot damage due to electroporation can be an issue and the
use of chemical delivery methods can restrict the range of
available surface chemistries.
[0170] Example 2 has also demonstrated the device's ability to
deliver proteins to the cell cytosol. Data and modeling estimates
indicate that the current subject matter could deliver
10-100.times. more material per cell relative to previous
practices, such as the use of cell penetrating peptides or
electroporation. This improvement in delivery rates provides a
powerful method for use in protein-based cell reprogramming, for
example, where the delivery of transcription factors to the cell
cytosol is a major hurdle to developing reliable methods of iPSC
generation. One may also use the current subject matter to study
disease mechanism by delivering various proteins/peptides of
interest. Indeed the current subject matter can be used for high
throughput screening of peptide libraries because, unlike most CPP
or nanoparticle-based techniques, the current subject matter is
insensitive to protein structure and chemistry, does not rely on
endocytotic pathways, and does not affect protein
functionality.
[0171] Because the current subject matter has demonstrated the
potential for delivery to primary cells, cytosolic delivery by
rapid mechanical deformation can be implemented as an ex-vivo
treatment mechanism. In this approach, the patient's target cells,
isolated from the blood or other tissue, are treated by the device
outside of the patient's body and then re-introduced into the body.
Such an approach takes advantage of the increased delivery
efficiency of protein or nanoparticle therapeutics and is safer
than existing techniques because, it obviates the need for
potentially toxic vector particles and mitigates any potential
side-effects associated with Reticuloendothelial clearance and
off-target delivery.
[0172] In example 2, the silicon-based devices were fabricated at a
microfabrication facility using photolithography and deep reactive
ion etching techniques. In this process, 6'' silicon wafers with a
450 .mu.m thickness are treated with Hexamethyldisilazane (HMDS),
spin coated with photoresist (OCG934, FujiFilm) for 60 s at 3000
rpm, exposed to UV light (EV1-EVG) through a chrome mask with the
constriction channel design, and developed in AZ405 (AZ Electronic
Materials) solution for 100 s. After 20 min of baking at 90.degree.
C., the wafer was etched by deep reactive ion etching (SPTS
Technologies) to the desired depth (typically, in this example, 15
.mu.m). Piranha treatment (H2O2 and H2SO4) were used to remove any
remaining photoresist after the etching process was complete. To
etch the access holes (i.e. inlet and outlet) the process was
repeated on the opposite side of the wafer (i.e. the one not
containing the etched channels) using a different mask, which
contains the access hole patterns, and a thicker photoresist AZ9260
(AZ Electronic Materials).
[0173] Oxygen plasma and RCA cleaning were used to remove any
remaining impurities. Wet oxidation was used to grow 100-200 nm of
silicon oxide before the wafer was anodically bonded to a Pyrex
wafer and diced into individual devices. Each device was
individually inspected for defects prior to use.
[0174] Before each experiment, devices were mounted onto a holder
with inlet and outlet reservoirs (all custom designed and produced
by Firstcut). These reservoirs interface with the device using
Buna-N O-rings (McMaster-Carr) to provide proper sealing. The inlet
reservoir is connected to a pressure regulator system using Teflon
tubing to provide the necessary driving force to push material
through the device. Example 2 can accommodate pressures up to 70
psi.
[0175] Example 2 cell culture included HeLa (ATCC), GFP expressing
HeLa, and DC2.4 (ATCC) cells were cultured in high glucose
Dubelco's modified essential medium (DMEM, Mediatech) supplemented
with 10% fetal bovine serum (FBS) (Atlanta Biologics) and 1%
Penicillin Streptomycin (Mediatech). Primary human fibroblast cells
(NuFF) (Globalstem) were cultured in high glucose DMEM supplemented
with 15% FBS. Cells were kept in an incubator at 37.degree. C. and
5% CO.sub.2. When applicable, adherent cells were suspended by
treatment with 0.05% Trypsin/EDTA (Mediatech) for 5-10 min.
[0176] Mouse embryonic stem cells (mESC) were grown on mouse
embryonic fibroblasts (Chemicon) in media consisting of 85% knock
out DMEM, 15% fetal bovine serum, 1 mM glutamine, 0.1 mM beta
mercaptoethanol, and 1% non-essential amino acids supplanted with
1000 units/mL LIF (Millipore, USA). Cells were passaged every 2-3
days using 0.25% Trypsin/EDTA. When treated with the device, the
mESCs were able to re-form colonies and retained normal morphology
even 2 weeks after treatment.
[0177] To perform an experiment for example 2, cells were first
suspended in the desired delivery buffer (growth medium, phosphate
buffered saline (PBS), or PBS supplemented with 3% FBS and 1% F-68
Pluronics (Sigma)), mixed with the desired delivery material and
placed in the device's inlet reservoir. This reservoir is connected
to a compressed air line controlled by a regulator and the selected
pressure (0-70 psi) was used to drive the fluid through the device.
Treated cells are then collected from the outlet reservoir. Cells
were incubated at room temperature in the delivery solution for
5-20 min post-treatment to ensure pore closure before being subject
to any further treatment.
[0178] In example 2 experiments comparing delivery of different
sized dextrans or protein vs. dextran, the molecules of interest
were co-delivered i.e. they were used in the same experiment, with
the same cell population, on the same device and differentiated
based on their fluorescent labels. All experimental conditions were
carried out in triplicate and the error bars represent two standard
deviations.
[0179] To deliver fluorescently labeled dextran molecules
(Invitrogen) or Apolioprotein E (Invitrogen), the example 2
experiments were conducted as described above such that the
delivery buffer contained 0.1-0.3 mg/ml of dextran or 1 mg/ml of
Apolioprotein E respectively.
[0180] To deliver fluorescein conjugated BSA (Invitrogen), cells
were first incubated in culture media containing 5 mg/ml unlabeled
BSA (Sigma) for 2 hours at 37.degree. C. and then treated with the
example 2 device using a delivery buffer containing 1 mg/ml of the
fluorescein conjugated BSA. The pre-incubation step was intended to
minimize non-specific binding of fluorescently labeled BSA to the
cell surface.
[0181] GFP knockdown for example 2 was measured as the percentage
reduction in a cell population's average fluorescence intensity
relative to untreated controls. Lipofectamine 2000+siRNA particles
were prepared by combining 1 .mu.g of siRNA with 1 .mu.l of
Lipofectamine 2000 reagent in 1000 of PBS. After 20 min of
incubation at room temperature, 20 .mu.l of this mixture was added
to each experimental well containing .about.20,000 cells and 100
.mu.l of media. The cells were allowed to incubate with the
particles for 18 hours prior to analysis.
[0182] In example 2, gold nanoparticles were prepared by
conjugating thiol terminated, 1000 MW polyethylene glycol (PEG) to
the nanoparticle surface, excess PEG was then washed four times by
centrifugation (10,000 rcf for 30 min) and the resulting material
suspended in PBS to a final concentration of 100 nM. To image GNP
delivery to HeLa cells, the cells were suspended in PBS
supplemented with 3% FBS, 1% F-68 Pluronics and 47 nM of GNP;
treated by a 10 .mu.m-6 .mu.m.times.5 device and fixed in 2.5%
(w/v) glutaraldehyde, 3% (w/v) paraformaldehyde, and 5.0% (w/v)
sucrose in 0.1M sodium cacodylate buffer (pH 7.4). After an
overnight fixation, the cells were post-fixed in 1% (w/v) OsO4 in
veronal-acetate buffer for 1 h. They were then stained en bloc
overnight with 0.5% uranyl acetate in veronal-acetate buffer (pH
6.0), dehydrated, and embedded in Spurr's resin. Sections were cut
on a Reichert Ultracut E (Leica) at a thickness of 70 nm with a
diamond knife. Sections were examined with an EM410 electron
microscope (Phillips).
[0183] In example 2, a Neon electroporation system (Invitrogen) was
used to transfect NuFF cells with fluorescein labeled 70 kDa and
pacific blue labeled 3 kDa dextrans. Manufacturer's procedure was
followed in washing cells and suspending them in the appropriate
buffers. Cells were treated using a 100 tip at a density of
10.sup.7 cells/ml with a dextran concentration of 0.1 mg/ml. The
three conditions used were as follows: 1) One 20 ms pulse of 1700V
2) Three 10 ms pulses of 1600V 3) Two 20 ms pulses of 1400V.
Condition 1 and 3 were both recommended by the manufacturer as the
optimal conditions for transfection of human fibroblast cells with
eGFP plasmid delivery efficiencies of 84% and 82% respectively.
[0184] In example 2 confocal images, samples were centrifuged at
800 rcf for 4 min and washed 2-3 times with PBS prior to imaging.
Confocal images were taken on live cells using the C1 confocal
add-on unit on a Nikon TE2000-U inverted microscope with a
60.times. water-immersion lens. Fluorescence samples were excited
by a 405 nm laser and detected using a standard DAPI filter
(Nikon).
[0185] In example 2 fluorescence microscopy, samples were
centrifuged at 800rcf for 4 min and washed 2-3 times with PBS prior
to imaging. Images were obtained using an Axiovert 200 (Zeiss)
inverted microscope equipped with Neofluar lenses (Zeiss).
Fluorescence excitation was provided by a X-cite 120Q mercury lamp
(Lumen Dynamics). The microscope is fitted with a Hamamatsu
C4742-95 camera (Hamamatsu) and images were analyzed by ImageJ
(NIH).
[0186] In example 2 flow cytometry, for analysis of cells after a
delivery experiment, cells were washed 2-3 times with PBS (>100
.mu.l per well in a 96 well plate). These were then re-suspended in
PBS supplemented with 3% FBS, 1% F-68 Pluronics and 10 ug/ml
propidium iodide (Sigma). Cells were analyzed on an LSR Fortessa
(BD Biosciences) or FACSCanto (BD Biosciences) equipped with a high
throughput sampling robot. The 405 nm and 488 nm lasers were used
for the excitation of the desired fluorophores. Propidium iodide
(live/dead stain), fluorescein and pacific blue signals were
detected using 695 nm long pass, 530/30 and 450/50 filters
respectively. Data analysis was conducted using FACS Diva (BD
Biosciences) and FlowJo (FlowJo) software.
Example 3--Stem Cells and Immune Cells
[0187] Proteins, nanoparticles, siRNA, DNA and carbon nanotubes
were successfully delivered to eleven different cell types,
including embryonic stem cells and immune cells. Indeed, the
ability to deliver structurally diverse materials and its
applicability to difficult to transfect primary cells indicate that
the device and methods have wide applicability in research and
clinical applications.
[0188] In example 3, each device consists of 45 identical, parallel
microfluidic channels, containing one or more constrictions, etched
onto a silicon chip and sealed by a Pyrex layer. The width and
length of each constriction (described in more detail below) range
from 4-8 .mu.m and 10-40 .mu.m respectively. The example 3 device
was typically operated at a throughput rate of 20,000 cells/s,
yielding close to one million treated cells per device prior to
failure, due to clogging. The parallel channel design was chosen to
increase throughput, while insuring uniform treatment of cells,
because any clogging or defects in one channel cannot affect the
flow speed in neighboring channels (the device can be operated at
constant pressure). Prior to use, the device can be first connected
to a steel interface that connects the inlet and outlet reservoirs
to the silicon device. A mixture of cells and the desired delivery
material can be then placed into the inlet reservoir and Teflon
tubing is attached at the inlet. A pressure regulator can be then
used to adjust the pressure at the inlet reservoir and drive the
cells through the device. Treated cells can be collected from the
outlet reservoir.
[0189] Parameters that influence delivery efficiency that have been
identified (see, e.g., example 2 above) can include cell speed,
constriction dimensions and number of constrictions (thereby
altering the shear and compression rates experienced by the cells).
For example, delivery efficiency of membrane impermeable, pacific
blue labeled 3 kDa dextran molecules to live HeLa cells increases
monotonically with cell speed across different constriction designs
(e.g., FIG. 25 at 2500). Constriction dimensions also impact
delivery; increasing the constriction length from 20 .mu.m to 40
.mu.m almost doubled delivery efficiency at all operating speeds
(e.g., FIG. 25 at 2500), with minimal effect on viability (e.g.,
FIG. 25 at 2510). Decreasing constriction width had a similar
effect. Increasing the number of constrictions in series also
increased delivery efficiency such that a device with five 10 .mu.m
length constrictions in series outperformed a single 10 .mu.m, 20
.mu.m or 40 .mu.m length design across all cell speeds (e.g., FIG.
25 at 2500 and 2510). In these data, the 0 mm/s data points
correspond to the control case whereby the cells undergo the same
treatment as the other samples but are not passed through the
device thus reflecting any endocytotic or surface binding
effects.
[0190] To investigate the versatility of the technique its ability
to deliver model dextran molecules to several cell types that are
traditionally difficult-to-transfect was assessed, especially
immune cells and stem cells. Fluorescently labeled 70 kDa and 3 kDa
dextran were used for these experiments because they are similar in
size to many protein and siRNA molecules respectively, easy to
detect by flow cytometry, and have minimal surface binding effects
as they are negatively charged.
[0191] FIG. 30 illustrates applicability of the current subject
matter across cell types. At 3000 is shown delivery efficiency and
viability of NuFF cells treated with a 30 .mu.m-6 .mu.m device to
deliver 3 kDa and 70 kDa dextran. At 3010 is shown delivery
efficiency and viability of spleen isolated, murine dendritic cells
treated with a 10 .mu.m-4 .mu.m device to deliver 3 kDa and 70 kDa
dextran. At 3020 is shown delivery efficiency and viability of
murine embryonic stem cells treated with a 10 .mu.m-6 .mu.m device
to deliver 3 kDa and 70 kDa dextran. At 3030 is shown delivery
efficiency of 3 kDa and at 3040 is shown 70 kDa dextran to B-cells
(CD 19.sup.+), T-cells (TCR-.beta..sup.+) and Macrophages (CD11b)
isolated from whole mouse blood by centrifugation and treated by 30
.mu.m-5 .mu.m and 30 .mu.m-5 .mu.m.times.5 devices at 1000 mm/s. 3
kDa and 70 kDa dextran were labeled with pacific blue and
fluorescein respectively. All data points were run in triplicate
and error bars represent two standard deviations.
[0192] Using various device designs dextran molecules were
delivered to newborn human foreskin fibroblasts (NuFF) (3000),
primary murine dendritic cells (3010), and embryonic stem cells
(3020). These experiments yielded minimal loss (<25%) in cell
viability (3000, 3010, and 3020) and results in murine embryonic
stem cells indicate that the method does not induce
differentiation. In further studies, white blood cells (buffy coat
layer) were isolated from murine blood by centrifugation and
treated them with the device. B cells, T cells and Macrophages, as
differentiated by antibody staining, indicated successful delivery
of both 3 kDa and 70 kDa dextran (3030 and 3040).
[0193] In order to illustrate the current subject matter's
potential in addressing current delivery challenges, a number of
experiments were conducted in possible applications ranging from
cell reprogramming to carbon nanotube based sensing. In addition to
the application specific materials detailed below, the current
subject matter has demonstrated the successful delivery of a variet
of test payloads such as Apolioprotein E, bovine serum albumin and
GFP-plasmids.
[0194] FIG. 31 illustrates data from nanomaterial and antibody
delivery. At 3100 is shown delivery efficiency and viability of
HeLa cells treated with a 10 .mu.m-6 .mu.m.times.5 device to
deliver pacific blue labeled 3 kDa dextran and Cy5 labeled, DNA
wrapped, carbon nanotubes. At 3110 is shown bright-field cell
images overlaid with Raman scattering in the G-band (red) to
indicate delivery of carbon nanotubes in treated cells (left) vs.
endocytosis (right). Scale bars at 2 .mu.m. At 3120 is shown
fluorescent micrograph of a HeLa cell 18h after delivery of pacific
blue labeled 3 kDa dextran (middle panel) and antibodies to tubulin
with an Alexa Fluor 488 tag (right panel). Scale bars at 3 .mu.m.
At 3130 is shown delivery efficiency and viability of HeLa cells
treated with a 10 .mu.m-6 .mu.m.times.5 device, at 500 mm/s, to
deliver Alexa Fluor 488 labeled anti-tubulin antibodies. Delivery
efficiency at different antibody concentrations is compared to an
endocytosis control at 100 .mu.g/ml and untreated cells
[0195] Verification of successful delivery of carbon nanotubes
(encapsulated by a DNA oligonucleotide) by flow cytometry (3100)
and Raman spectroscopy (3110). Antibodies to tubulin were also
delivered (3120 and 3130) using this technique, yielding a diffuse
distribution throughout the cell that would be consistent with
cytosolic delivery. The aforementioned materials are currently
difficult to deliver to the cell cytosol and each material often
requires a specialized modification to facilitate delivery. In
example 3, all four materials were delivered to HeLa cells using
the same set of conditions on a 10 .mu.m-6 .mu.m.times.5
device.
[0196] Efficient delivery of proteins to primary cells can enable
several therapeutic applications. A challenge in cell
reprogramming, for example, is the inefficiency of previous CPP
based protein delivery methods. FIG. 32 illustrates protein
delivery applications. At 3200 is shown a western blot analysis of
c-Myc, Klf-4, Oct-4 and Sox-2 delivery by cell penetrating peptides
versus a 10 .mu.m-6 .mu.m device to NuFF cells. Each of the four
proteins has an additional 9 arginine (9R) groups to facilitate
uptake. The lysate (Ly) columns correspond to the protein content
of cells that are washed and lysed while the soup columns
correspond to the protein content of the media environment. At 3210
is shown delivery efficiency and viability of spleen isolated,
dendritic cells treated with a 10 .mu.m-4 .mu.m device to deliver
pacific blue labeled 3 kDa dextran and Alexa Fluor 488 labeled
ovalbumin. All data points were run in triplicate and error bars
represent two standard deviations.
[0197] The ability to deliver four exemplary transcription factors
(Oct4, Sox2, c-Myc, and Klf-4) to human fibroblast cells were
examined and compared to a CPP method (3200). The results show that
in addition to not relying on endocytosis, which can leave much
material trapped in endosomes, delivery by rapid mechanical
deformation yields significantly higher delivery efficiency for all
4 proteins. This result is in line with the aforementioned
simulation work which indicated the current subject matter can have
a 10-100.times. improvement in protein delivery relative to
CPPs.
[0198] Antigen presentation in dendritic cells (DCs) is another
area in which the current subject matter offers an advantage.
Researchers have been exploring methods to express antigens on the
MHC class I receptors of DCs so as to induce a potent cytotoxic T
cell response. The current subject matter, which has direct
clinical implications in preparation of a cancer vaccine, for
example, relies on the cytosolic delivery of antigenic proteins,
because MHC class I presentation is almost exclusive to cytosolic
proteins. By facilitating direct cytosolic delivery of material,
the current subject matter serves as a platform for generating an
in vivo cytotoxic T cell response against a specific antigen. To
illustrate this capability, Alexa 488 labeled ovalbumin, a model
antigen protein, was successfully delivered to murine dendritic
cells derived from the spleen (3210). Despite the higher rate of
endocytosis in this cell type, the device produced a significant
increase in delivery rates relative to the endocytosis controls (0
mm/s). Moreover, the delivered material is present in the cytoplasm
by virtue of the cell deforming mechanism. This feature is
particularly important for antigen delivery, bevause cytoplasmic
pressure is a critical requirement for MHC class I antigen
presentation.
[0199] The system of example 3 is an enabling research tool with
its ability to deliver carbon nanotubes, gold nanoparticles and
antibodies (FIG. 31)--three materials that are difficult to deliver
with current techniques. The current subject matter significantly
expands ability to probe intracellular processes by facilitating
antibody and quantum dot staining of live cell structures/proteins
and enabling the use of carbon nanotubes as a cytosolic molecular
probe or chemical sensor. As a robust method of protein delivery,
it can be used for high throughput screening of peptide/protein
libraries because, unlike most CPP or nanoparticle-based
techniques, this method is insensitive to protein structure and
chemistry, does not rely on endocytotic pathways, and does not
affect protein functionality.
[0200] Additionally, the current subject matter is useful for
therapy (FIG. 32). For example, a patient's target cells are
isolated from the blood or other tissue, treated by the device to
deliver the desired therapeutic, and re-introduced into the body.
Such an approach capitalizes on increased delivery efficiency of
therapeutic macromolecules and is safer than existing techniques,
because it obviates the need for potentially toxic vector particles
and mitigates any potential side-effects associated with
Reticuloendothelial clearance and off-target delivery.
Example 4--Personalized Cancer Vaccinations
[0201] Current challenges in the intracellular delivery of
macromolecules are a significant barrier to better understanding
disease mechanisms and the implementation of novel therapeutic
approaches. Despite recent advances in delivery technology,
treatment of patient-derived cells remains a challenge and current
methods often rely on toxic electrical fields or exogenous
materials. The microfluidic platform and related systems and
methods rely on the mechanical deformation of cells to facilitate
delivery. This controlled, physical approach has produces results
in previously challenging areas, such as protein-based cell
reprogramming and quantum dot delivery.
[0202] The most effective and direct way to influence the behavior
of a cell is by delivering active agents to the cell cytoplasm.
Intracellular delivery of macromolecules thus plays a critical role
in research and development (R&D), with applications ranging
from drug discovery to the study of biochemical processes to
therapeutic applications. Current methods, however, have
limitations. They often have low efficacy in patient-derived
(primary) cells, rely on toxic electrical fields or exogenous
material, and are ill-suited for the delivery of structurally
diverse materials, such as proteins.
[0203] A robust delivery platform capable of addressing these
issues enables significant advances in biological research and
serves as the basis for a new generation of therapies such as
personalized cancer vaccination. An effective protein delivery
method for immune cells, for example, can serve as a cancer
vaccination platform.
[0204] As described above, the microfluidic devices, related
systems and methods described herein facilitates intracellular
delivery of material by rapidly deforming a cell as it passes
through a constriction. The deformation process causes transient
disruption of the cell membrane and thereby enables the passive
diffusion of material from the surrounding buffer into the cell
cytosol. By eliminating the need for the exogenous materials and
electrical fields that current methods rely on, this approach
provides a simplified, robust approach to delivery with reduced
toxicity. Hence, this method can serve as a broad platform to
intracellular delivery of macromolecules with advantages in some
research and clinical applications, such as cancer vaccination.
[0205] FIG. 34 is an illustration depicting a system in which a
patient's blood is treated by a microfluidic device for the
delivery of macromolecules. One embodiment of the current subject
matter includes a system in which dendritic cells (DCs), isolated
from a patient's blood, are treated by the device, ex vivo, to
activate them against a particular cancer antigen and then
reintroduced into the patient's blood stream. For example,
delivered antigen is a commonly expressed protein known to be
associated with a particular disease or a patient-specific one
obtained from a biopsy. By delivering cancer antigens directly to
the DC cytoplasm, one can exploit the MHC-I antigen presentation
pathway and induce a powerful cytotoxic T lymphocyte (CTL) response
in the patient. These activated T-cells then seek out and destroy
any cancerous cells which express the target antigen. The
platform's flexibility in using established, disease-specific
antigens or those derived directly from a patient's tumor allow it
to treat patients that are resistant to other therapies. Indeed
this provides a personalized, targeted disease response with
minimal side-effects. This embodiment can be implemented in a
typical hospital laboratory (<1 hr per treatment) with a trained
technician. Due to its small size and relative simplicity, a
patient operated treatment system can also be used.
[0206] The cancer vaccine method has been demonstrated in a mouse
model. The system has been used for the successful delivery and
processing of ovalbumin, a model antigen, to murine dendritic cells
and as indicated by increased presentation of the antigen, SIINFEKL
peptide, on MHC class I receptors. These treated dendritic cells
promote a proliferative cytotoxic T lymphocyte (CTL) response in
vitro. Treated DCs are reintroduced into the animal to generate an
in vivo CTL response. The devices are useful to deliver antigens to
DCs isolated from human blood.
[0207] The data indicates that the device is capable of delivering
material to sensitive, primary cells (including DCs) without
causing excessive cell death. Thus, cell damage is not a serious
issue. Immune response can be improved by increasing the number of
treated cells, increasing the quantity and diversity of antigen
delivered, and/or co-delivering activating factors, such as
lipopolysaccharide. Little or no toxicity has been observed with
cells treated with the device, thus making the methods not only
feasible but advantageous for therapeutic human and veterinary
applications.
Example 5--Blood Cancer Treatment
[0208] As described in example 4, rapid mechanical deformation of
cells can provide a robust means of delivering antigens to
dendritic cells (DCs) and thus can be a platform for cellular
therapies. This system is based on the discovery that rapid
mechanical deformation of cells can cause the formation of
transient membrane pores that enable diffusive delivery of material
from the surrounding medium. Unlike existing therapies discussed
earlier, this method does not rely on custom fusion proteins,
antigen cross-presentation, viral vectors, nanoparticles or
endocytosis mechanisms; therefore, it provides great improvements
in efficacy in vivo while reducing therapeutic costs. The
flexibility and simplicity of this fundamentally different approach
enables a broad platform for dendritic cell activation that can
induce a CD8 response against a variety of cancer antigens. The
system can target blood cancers, such as B cell lymphoma, which are
more amicable to immune therapies, as well as several additional
cancer types (e.g. Melanoma, Pancreatic cancer, etc.) and provides
a vital, personalized new approach to combat the disease.
[0209] Cellular therapy against cancer is an attractive option due
to its ability to activate the patient's immune system and drive a
long-lasting antigen-specific CD8 T cell response against the
disease. These therapies, such as the recently approved
Provenge.RTM. for prostate cancer, have minimal side-effects
relative to chemotherapies and radiation treatment. However, one of
the greatest barriers to developing cellular therapies has been
achieving proper antigen presentation by delivering antigens into
the cell cytoplasm. Traditionally, activation of a CD8 effector
response differentiates from a CD4 response by the location of the
foreign protein entering the antigen presenting cell (e.g.
dendritic cell). Proteins found in the cytoplasm induce a CD8
response while extracellular proteins captured by endocytosis
induce a CD4 response. Since the mechanism of cross presentation
within antigen presenting cells remains elusive, one must develop a
reliable method of delivering the desired antigen directly to the
cell cytoplasm to advance therapies utilizing the potent cytotoxic
effector CD8 response. A robust, effective method of cytoplasmic
delivery to dendritic cells is used as a platform to induce immune
responses against a variety of cancer types.
[0210] Experiments on HeLa cells, illustrated in FIG. 8A and FIG.
11, have indicated that rapid mechanical deformation of the cell
results in the formation of transient injuries/pores in the cell
membrane, which enable the passive diffusion of material from the
surrounding buffer into the cell cytoplasm. This previously
unreported phenomenon produces cells that are viable and
proliferate normally after treatment. Experiments have indicated
that larger molecules exhibit lower delivery rates than smaller
ones, thereby indicating a diffusive mechanism. Successful siRNA
delivery and diffuse cytosol staining, as measured by confocal
microscopy, also indicate that the delivered materials are in the
cytosol and in an active/accessible state. Traditionally, methods
of antigen delivery often show promise in cell lines but fail to
translate to primary immune cells. The delivery method described
herein, however, is independent of endocytotic pathways or cellular
response to exogenous materials, which can vary significantly
across cell types, and relies primarily on membrane bilayer
properties. Hence, due to its simplicity and novel pathway, this
technology is more amicable to transitioning from cell line to
primary immune cell delivery and thus provides a major improvement
in antigen presentation.
[0211] MHC class I presentation of antigens and dendritic cell
maturation can be analyzed by antibody staining in response to
delivery of ovalbumin protein. T cells harvested from OT-I and
OT-II TCR transgenic mice can also be used to measure CD8 and CD4
proliferation in response to ovalbumin and/or SIINFEKL antigen
loading. The system can be optimized for increased CD8
proliferation as compared to dendritic cells primed by endocytosis
alone.
[0212] Primary murine dendritic cells are purified by MACS
CD11c+separation (Miltenyi Biotec, Germany) from the spleens of B6
mice. A device that contains constrictions with a 6 .mu.m channel
width capable of porating a 13 um HeLa cell's membrane can be used.
Due to the smaller size of these dendritic cells, microfabrication
and testing devices with channel widths of 3-5 .mu.m can be
utilized performance. Existing protocols for photolithography and
deep reactive ion etching can be modified to enable efficient
manufacture of these devices. Fluorescently labeled dextran
molecules can be used as model molecules to assess delivery
efficiency by FACS. Subsequently, ovalbumin protein can be
delivered to the cells and assayed by a western blot to confirm
protein uptake in primary cells.
[0213] Dendritic cell maturation can be examined by CD80 and CD86
antibody staining to show the rapid deformation method induces cell
maturation. The use of extracellular TLR agonists, such as
lipopolysaccharide (LPS), can be considered if it is deemed
necessary to manually induce DC maturation post-delivery. Ovalbumin
protein can be delivered to dendritic cells and antigen
presentation can be quantified by MHC-I SIINFEKL antibodies.
Additionally, antigen presentation efficiency in response to
delivery of TCR specific peptides can be assessed to show the
system's ability to deliver/present antigenic proteins vs.
peptides. Subsequently, CD8 and CD4 T cells can be harvested from
TCR transgenic OT-I and OT-II mice respectively, stained by CFSE,
and co-cultured with ovalbumin treated dendritic cells for 5 days.
T cell proliferation of both subsets can be measured by FACS.
Device design can be optimized to produce increased levels of
functional CD8 T cell populations in comparison to conventional in
vitro methods (e.g. endocytosis).
[0214] A versatile ability to induce antigen-specific CD8 responses
has been a goal of cancer cellular therapies that has proved
elusive thus far due to inefficient antigen presentation or
inadequate flexibility of the delivery method. Existing methods
have a number of drawbacks including their reliance on damaging
electrical fields, the use of exogenous materials, protein sequence
modification and/or endocytotic pathways to facilitate antigen
delivery. The current subject matter, however, provides a
fundamentally different approach to cytosolic delivery that does
not suffer from any of the aforementioned problems. Moreover, by
nature of its poration-diffusion mechanism, this method is broadly
applicable across antigen types and could thus address a range of
target cancers. The same mechanism can even be used to introduce
additional signaling molecules to improve DC maturation/activation
to produce a more potent T cell response. Such a broad-based
platform is more versatile and robust than any existing antigen
delivery/presentation mechanisms under investigation for cancer
vaccines.
[0215] Example 5 can have a broad impact. Given the immense social
burden of cancer across the country (estimated 570,000 deaths in
the U.S. in 2011); cancer is likely to afflict a significant
proportion of the population The afflicted population benefit from
the development of novel cell therapies that harness the power of
the patient's immune system to combat the disease. The current
subject matter is a more efficient, personalized treatment platform
for a variety of cancer types, such as blood cancers, e.g,
leukemias, lymphomas, and multiple myelomas, as well as
myeloproliferative neoplasms and myelodysplastic syndromes. The
methods are also particularly useful for treatment of metastatic
cancers, e.g., due to their propensity to disseminate via blood
circulation. Cancers with unknown antigen epitope, for example, may
be treated by digesting a tumor biopsy sample, delivering the
lysate to the patient's DCs, and reintroducing the DCs into the
body. This would enable one to activate the host's T cells against
a broad range of cancer antigens thereby ensuring effective,
multi-target treatment. This personalized aspect could be of
particular interest to people who could develop rare forms of
cancer, often ill-served by current treatments, due to exposure in
hostile environments. By tailoring treatment to the individual's
disease, this method can provide timely, effective care in even the
most aggressive cases, such as multi-drug resistant cancers. This
immune-based therapy can also be particularly effective at
preventing metastasis (responsible for .about.90% of cancer related
deaths) as CD8 T cells may easily locate and destroy metastatic
cells while the immunological memory provided through these
treatments could prevent future relapse. In addition, as a research
tool, this method can enable unprecedented mechanistic studies of
antigen processing to better understand the process of antigen
cross-presentation and hence improve the efficacy of
existing/alternative immune activation methods.
Example 6--Cell Reprogramming
[0216] Stem cells play a critical role in current research in
regenerative medicine, especially within the rapidly expanding
field of tissue engineering. iPSCs are of particular interest due
to their capacity for self-renewal, demonstrated ability to
differentiate into any cell type, and autologous (patient specific)
characteristics. Thus, iPSCs provide an opportunity to derive
multi-lineage progenitor cells from a common pluripotent source,
which may be combined into distinct yet interactive tissue
compartments. Moreover, these cells could eventually obviate the
need for human embryonic stem cells (hESCs) in clinical
applications thus avoiding many of the moral and ethical debates
that have plagued these cell types. Furthermore, patient-derived
iPSCs avoid or minimize the immune rejection problems of
hESC-derived cells. Thus, current research is largely focused on
devising efficient, virus-free, protocols to produce large numbers
of iPSCs.
[0217] iPSCs were originally generated by reprogramming adult
murine and human fibroblasts (HFs) to a pluripotent state based on
retro-viral overexpression of the 4 transcription factors Oct 3/4,
Sox2, c-Myc and Klf4. These iPSCs are not only largely identical to
ES cells in global gene expression, DNA methylation, and histone
modification, but are also able to differentiate into cell types
representing all 3 germ layers. While iPSC technology has enormous
potential for biomedical research and cell-based therapy, major
obstacles must be overcome to realize its full potential. For
instance, most iPSC lines have been derived from various somatic
cells by retroviral or lentiviral introduction of reprogramming
factor-encoding genes, resulting in multiple chromosomal
disruptions by viral vector integration, any of which may cause
genetic dysfunction and/or tumor. In addition, reprogramming
transgenes (in particular, c-Myc and Klf4) are closely associated
with oncogenesis raising the possibility that its residual
expression and/or reactivation may cause tumor formation. Thus,
many laboratories recently explored different genome
non-integrating approaches such as adenoviruses, episomal vectors,
mRNAs, and microRNAs. Notably, it has been shown that iPSCs can be
generated by direct delivery of the four reprogramming factors (Oct
3/4, Sox2, c-Myc and Klf4) fused to cell-penetrating peptides
(CPP). While it has been reported that generation of mouse iPSCs
can occur by delivery of four CPP-fused factors expressed in E.
coli, it can be shown that human iPSCs can be generated by four
CPP-fused factors expressed in mammalian cells. However, both
studies reported that the reprogramming efficiencies of
protein-based reprogramming is very low (<0.01%). Since
protein-based reprogramming does not involve any type of genetic
material (DNA or RNA) and vector vehicle (virus or plasmid), direct
delivery of proteins provides one of safest reprogramming
procedures. It has been shown that protein-based human iPSCs
efficiently generated functional dopamine neurons without abnormal
properties associated with viral genome integration. Since the
efficiency of protein-based reprogramming can be improved by the
current subject matter using the delivery platform technology, it
widely opens the possibility to generate clinically viable iPS
cells. Moreover, this approach enables a finer level of control
over cellular function by circumventing the stochastic processes
that govern translation and/or transcription in mRNA, plasmid and
viral reprogramming. Direct protein delivery thus provides two
fundamental advantages over alternative methods by obviating the
risk of mutagenic insertion and enabling more accurate control of
the highly sensitive reprogramming process. The delivery technology
described herein has demonstrated its ability to deliver proteins
at high efficiencies to HFs and stem cells. Experiments comparing
this technique's delivery capabilities to existing cell penetrating
peptide methodologies have shown a significant increase in delivery
using this approach (potentially 100.times. higher based on
simulations). Moreover, its physical poration mechanism eliminates
the need for chemical modification or the use of exogenous
compounds that are involved in alternative protein delivery
methods. Small molecules, siRNA and other factors can also be
co-delivered during reprogramming as the method is agnostic to the
type of material being delivered. This system thus provides a
unique tool for inducing cell reprogramming through direct protein
delivery. This simple mechanism of action (i.e. diffusion through
pores) also enables one to potentially predict and control delivery
quantities with high accuracy, thus facilitating optimization
studies to improve the understanding of reprogramming dynamics and
thereby greatly increasing efficiencies. Finally, one can deploy
this microfluidic technique as a medical device to generate iPSCs
for clinical tissue engineering and cell therapy applications.
[0218] Moreover, the applications of this system are not confined
to protein delivery. This technique can include into a universal
delivery method capable of delivering a range of macromolecules
(DNA, RNA, proteins, sugars, and peptides) to almost any cell type.
This enables a host of applications that are underserved by current
technologies. Current liposomal, nanoparticle, and
electroporation-based methods, for example, often struggle to
transfect certain primary cells (such as immune cells or stem
cells) and can be ineffective at delivering proteins and
nanoparticles (such as quantum dots). Peptide delivery for
therapeutic screening and disease mechanism applications can also
be addressed by this novel method whereas contemporary practices
often require chemical modification or encapsulation. One can also
use this method for nanoparticle-based sensing applications to
deliver modified quantum dots for organelle labeling and
mechanistic disease studies.
[0219] Intracellular delivery is a cornerstone of many biological
research applications ranging from fundamental studies of gene
expression, to disease mechanisms and, as addressed in this
application, generation of iPSCs. Established delivery methods,
such as liposomes, polymeric nanoparticles and electroporation
often involve the use of exogenous compounds as a delivery vehicle
(or electric fields in the case of electroporation) and are
material and/or cell specific. For example, lipofectamine
(Invitrogen) can deliver DNA and RNA molecules (to subsets of cell
lines or primary cells) but cannot form the proper complex to
deliver proteins or other macromolecules. Electroporation, on the
other hand, although promising in its ability to target a variety
of cell types, causes damage to the cell due to the high electric
fields and has had limited success in protein delivery. This makes
it particularly unsuitable for the multiple transfections required
in iPSC generation, for example. Membrane penetrating peptides are
another delivery technique that is largely specific to proteins.
These peptide-based methods, however, have unpredictable effects on
protein functionality and suffer from significant protein
degradation in the endosome. Hence, the current subject matter
describing a universal method capable of delivering a range of
macromolecules (DNA, RNA, proteins, peptides, small molecules),
with minimal cell death, enables unprecedented control over
cellular function in a single technology platform, thereby enabling
studies of disease mechanism, identification of macromolecular
therapeutic candidates, guided differentiation or reprogramming of
stem cells, and the development of diagnostic techniques with
reporter cell lines.
[0220] The microfluidic device described herein can serve as a
broad-based universal delivery platform. As a microfluidic device,
it enjoys the benefits of precise control over treatment conditions
on a single-cell level. The unique combination of single-cell level
control and macro-scale throughput places this device in a unique
position relative to existing delivery methods. Data thus far has
demonstrated the system's ability to deliver material to over 11
different cell types including cancer cell lines, embryonic stem
cells, primary fibroblasts, and primary lymphocytes. Its mechanical
poration mechanism has also enabled the delivery of previously
challenging materials such as carbon nanotubes and quantum
dots.
[0221] Previous work using recombinant proteins to produce iPSCs
have demonstrated prohibitively low efficiencies (<0.01%) and
are thereby unsuitable for wide-spread clinical application. The
device, systems, and methods mentioned herein, however, have
demonstrated their ability to deliver proteins directly to the
cytoplasm with high efficiency and minimal cell death thus
providing a compelling opportunity to produce substantial gains in
reprogramming efficiency through more effective delivery. By
directly determining the quantity of available protein one can
exercise accurate control over intracellular kinetics. Other
reprogramming methods (e.g., viral, plasmid and mRNA expression),
on the other hand, rely on stochastic effects to determine the
level of protein availability and are thus unsuitable for kinetic
studies. The low efficiency of current reprogramming methodologies
indicates that the process is highly sensitive to stochastic
variations and only a narrow range of transcription factor
expression levels will result in reprogramming. By directly
delivering proteins to the cytoplasm, one can exercise
unprecedented control over protein availability and thus more
consistently impose the exact conditions necessary for
reprogramming. These conditions, once identified and optimized, can
be reproduced accurately for every cell undergoing treatment and
thus dramatically improve reprogramming efficiency.
[0222] This technique enables/improves upon a variety of
intracellular delivery applications. In addition, the strictly
mechanical nature of the technique eliminates any potential
complications arising from the use of chemical agents or electric
fields. Data has not revealed any substantial changes in cell
behavior as a result of the treatment. Thus, this system is a
robust, high-throughput, high-efficiency, universal intracellular
delivery mechanism with particular utility in reprogramming
applications.
[0223] Evidence indicates that the rapid deformation that occurs as
a cell passes through the constriction induces the formation of
transient pores in the cellular membrane enabling diffusion of
macromolecules from the surrounding buffer into the cytosol. This
technique has been demonstrated in 11 different cell types
including cancer cell lines, primary fibroblasts, primary
lymphocytes, and embryonic stem cells (without causing
differentiation). One prototype is capable of treating
.about.20,000 cells/s and operating at a range of cell
concentrations (104-108 cells/ml). Issues pertaining to clogging
have also been largely mitigated by improving experimental
protocols and chip design such that each device is capable of
treating .about.1 million cells prior to clogging, with the option
of being cleaned and recycled. In addition, the multi-channel
design provides significant redundancy such that the clogging of
one channel does not affect the performance of the others. Pressure
driven flow (at controlled constant pressure) and the parallel
design of the channels ensure a consistent flow profile per channel
regardless of the percentage of clogged channels in the chip.
[0224] The device's ability to deliver dextran molecules to human
fibroblasts and embryonic stem cells has been demonstrated. FIG. 35
illustrates the potential advantages of cell reprogramming. At
3500, the delivery efficiency and viability of human embryonic stem
cells treated with a 10 .mu.m-6 .mu.m device to deliver 3 kDa
dextran. At 3510, a western blot analysis of c-Myc, Klf-4, Oct-4
and Sox-2 delivery by cell penetrating peptides versus a 10 .mu.m-6
.mu.m device to NuFF cells. The lysate (Ly) columns correspond to
the protein content of cells that are washed and lysed while the
soup columns correspond to the protein content of the media
environment. At 3520 confocal microscopy images of NuFF cells fixed
after delivery of the reprogramming factors. The proteins are
tagged using an Alexa 488 conjugated anti-FLAG antibody and the
nucleus is stained by DAPI.
[0225] Moreover, the devices delivery efficiency was compared to
that of a 9 arginine (CPP) method currently used for protein-based
reprogramming. The results (35 10) demonstrated a significant
increase in the quantity of c-Myc, Klf4, Oct4 and Sox2 delivered as
measured by western blot. Confocal microscopy then confirmed the
successful localization of these transcription factors to the cell
nucleus (3520). A simple 2-D diffusion model was developed in
COMSOL to simulate the delivery mechanism based on literature
values for particle diffusivities inside and outside the cell
cytoplasm. Fitting this model to the experimental data it can be
estimated that the technique delivers 10-40% of the delivery
material present in the buffer into the cell cytosol. By
comparison, CPP methods for protein delivery are estimated to
deliver only 0.1% of the buffer material to the cytosol. This
approach thus provides a robust increase in quantity of
reprogramming material delivered (10-100 fold). Moreover, it
ensures greater bioavailability of the delivered transcription
factors.
[0226] FIGS. 36A-36F depict the generation and characterization of
mouse and human iPSC lines by direct delivery of fused
reprogramming proteins. At 3600, starting mouse hepatocyte culture
(first image); morphology after 6 cycle protein treatments (second
image); established iPS colonies (third image); and AP staining of
established iPS colonies (fourth image). At 3610 immunostaining of
ESC markers (Nanog, Oct4 and SSEA1) in p-miPSC. Nuclei were stained
with DAPI (blue). At 3620 bisulfite sequencing analysis of the Oct4
promoter reveals almost complete epigenetic reprogramming in
p-miPSC-1 and p-miPSC-2 lines. Open and closed circles indicate
unmethylated and methylated CpG, respectively. At 3630, in vivo
differentiation potential was analyzed by injecting p-miPSCs into
immunodeficiency mice and by H&E staining of teratomas. The
resulting teratomas contained tissues representing all three germ
layers; ectoderm (neural tube or epidermis), mesoderm (cartilage or
muscle), and endoderm (respiratory epithelium or intestinal-like
epithelium) lineage cells. At 3640 chimeras derived from p-miPSC-1
(left panel) and p-miPSC-2 (right panel) at E13.5 fetuses show a
high level of GFP from injected p-miPSCs. At 3650 to 3670, human
iPSC lines, p-hiPSC-01 (3660) and p-hiPSC-02 (3670) are generated
by direct delivery of CPP-fused four reprogramming factors from
biopsied adult human fibroblasts (3650).
[0227] FIG. 37 depicts preliminary protein reprogramming results.
At 3700, a progression of morphological changes from fibroblasts
into colonies. White arrows indicate potential reprogrammed cells.
The red arrow points towards coalescing iPSCs forming a colony. At
3710 to 3760, expression of the human embryonic stem cell marker
Oct4, SSEA-4, Tra-60, Tra-80, Alkaline Phosphatase (AP) in iPSC
colonies. Were appropriate, the small box represents a DAPI counter
stain. Scale bars at 100 m.
[0228] Since protein-based human iPSCs were generated and
characterized, further fully reprogrammed mouse and human iPSCs
were generated by the previous delivery method of CPP-fused
reprogramming factors, as examined by all criteria including
epigenetic analyses, in vivo pluripotency, and chimera formation
(FIGS. 36A-36F). However, despite the use of partially purified
proteins, the reprogramming efficiency was still low (<0.1%) and
took longer than viral reprogramming methods. Thus, it was
attempted to use the device to deliver 4 reprogramming proteins
Oct4, Sox2, Klf4, and c-Myc to human fibroblasts at a buffer
concentration of 80 .mu.g/ml. Cells were treated 4 times with a 48
hr interval between each delivery. Following 14-20 days in culture
the first reprogrammed hiPSC-like colonies appeared. During this
time, it was observed that the transition in fibroblast morphology
as they formed iPSC colonies and they express several hESC markers
(FIG. 37).
[0229] Clathrin, caveolae and macropinocytosis are the three most
commonly proposed mechanisms for endocytotic internalization. To
examine whether endocytosis is involved in macromolecular delivery
following rapid cell deformation, it is possible to use known
chemicals to block these mechanisms. Specifically, chlorpromazine
can be used to inhibit clathrin mediated endocytosis; genisten to
inhibit caveolae mediated endocytosis; and 5-(N-ethyl-N-isopropyl)
amirolide (EIPA,) to inhibit macropinocytosis (all can be purchased
from Sigma Aldrich). HeLa cells can be incubated with
chlorpromazine (10 .mu.g/ml), genisten (200 .mu.M) and EIPA (25
.mu.M) for 2 hour prior to treatment. Dextran, dsRED and dsRED-9R
proteins can then be delivered by rapid deformation of the treated
cells. The respective delivery efficiencies, as measured by FACS,
can illustrate the influence of endocytosis inhibition on both CPP
and device-based delivery mechanisms. Co-localization experiments
with endosome markers (Invitrogen) using confocal microscopy can
also help determine the percentage of material that is sequestered
into endosomes.
[0230] It is possible to couple the rapid cell deformation system
with other established methods of delivery, such as
electroporation, to mitigate endocytotic mechanisms. Incorporating
electrodes near the constriction can couple deformation and
electroporation to enable delivery effects to yield enhanced system
performance relative to either individual method. In addition,
co-delivery of chemical agents such as Chloroquine (Sigma), various
polymers or endosome escape peptides, can be used to assist
endosome escape of delivered materials in the rapid cell
deformation system.
[0231] As a cell passes through the constriction, it experiences
brief (.about.10-100 us) but rapid shearing and compression.
Tangential shearing has been previously shown to induce pore
formation. However, the system also induces mechanical compression.
To evaluate these parameters, HeLa and HF cells can be incubated in
0.1 .mu.g/ml Lantrunculin A (Invitrogen) for 1 hour prior to
delivery to depolymerize the actin cytoskeleton. Fluorescently
labeled dextran (Invitrogen) can then be delivered to the treated
cell population using the rapid deformation device. These
experiments can also be repeated with cells that have been
incubated in 10 .mu.M Colchicine (Sigma) for 2 hours prior to
delivery to depolymerize the microtubule network. FACS analysis can
be used to measure delivery efficiencies of toxin treated cells
relative to untreated controls. Pore formation is believed to
correlate to the deformation rate of a cell in response to a given
geometry. The cytoskeleton's role in resisting deformation was
previously investigated using a device that probes cell
deformability to provide quantitative measurements of deformation
rates. In this method, electrodes are placed on either side of a
constriction and the change in capacitance between the two
electrodes is measured as a cell passes through. Changes in
capacitance across the constriction are then correlated to cell
transit time i.e. its deformation rate. The device's delivery
performance, in the experiments mentioned above, can be correlated
to these prior deformation studies to produce a quantitative
relationship between deformation rate and poration efficiency.
[0232] Similar to published studies characterizing sonoporation,
scanning electron microscopy (SEM) and transmission electron
microscopy (TEM) techniques can be employed on samples fixed at
defined time intervals after treatment to directly measure the size
and distribution of the proposed pores over time. Cell fixation can
be done at room temperature using a 25% Gluteraldehyde solution
(Sigma). The cell samples can then be dehydrated through successive
ethanol washes prior to imaging. Environmental SEM (ESEM)
techniques can be used to directly image fixed samples. Due to its
relatively low resolution (.about.200 nm), however, this technique
is suitable for detecting 1-0.5 .mu.m scale morphological changes
or injuries. Should ESEM fail in detecting any morphological
changes, the cells can be coated with a 1-10 nm layer of gold using
a vacuum evaporator to enhance resolution down to the nanometer
scale necessary to directly observe finer pore structures using
SEM. TEM can also be used as an alternative imaging technique
should SEM fail to produce the desired results. These techniques
enable one to distinguish between a local injury and uniform
poration mechanism of delivery and measure the average pore size
and distribution. Pore size and distribution in cells that
underwent rapid cell deformation can be compared to untreated
cells. A localized pore distribution on the membrane surface
indicates an injury model while a more uniform distribution
supports the uniform poration model.
[0233] COMSOL multiphysics software can be used to construct a 3D
model of the porated cell. Using published data on cytoplasm and
buffer diffusivities, combined with the appropriate pore models
from mechanistic studies it is possible to produce a predictive
model of delivery. The model emulates a porous membrane separating
a low diffusivity cytoplasm from a high diffusivity buffer region.
Under the model's assumptions the pores have a fixed size for a
fixed amount of time before instantaneous resealing. Dynamic pore
behavior, such as changes in shape and diameter, can be
incorporated into complex models through coupling with MatLab or
other software. Simulated predictions of delivery quantity can be
verified using experimental data based on FACS and gel
electrophoresis (e.g. western blots). These comparisons can be used
to fine-tune the model and hence enable it to predict the quantity
of material delivered. This model's predictive capabilities can
simulate the effects of varying pore size, pore opening time and
buffer concentrations and hence be used as a guide to future
studies.
[0234] Multi-physics simulations (e.g. COMSOL or CFD-ACE) can also
be used to model fluid flow throughout the device. These models can
be used to more accurately predict flow speeds and sheer stresses
in the inlet, outlet, and constriction regions. This data can be
used to elucidate links between sheer stress and flow speed to
delivery efficiency across constriction designs. Moreover, by
constructing a broad model of the device it is possible to study
the consistency of pressure drops between different channels and
adjust the inlet and outlet designs to ensure that all channels
operate under near-identical conditions so as to improve treatment
uniformity across the cell population.
[0235] The delivery phenomenon can be optimized primarily for
increasing delivery efficiency and cell viability. Population
uniformity (i.e. delivering a similar amount of material to each
cell) can be used as a secondary optimization parameter. Initial
results have identified cell speed, constriction length,
constriction width, and entry region shape as sensitive parameters.
Media composition, on the other hand, does not appear to be a major
factor. It is possible to construct a series of devices, which
systematically vary constriction length and width between 5-50
.mu.m and 4-8 .mu.m respectively. Different taper angles as the
main channel narrows to form the constriction are possible. The
experimental efficiency and viability data from these devices, as
measured by FACS, can be correlated to the aforementioned modeling
data to better understand the effects of sheer stress and
constriction dimensions. This process can be repeated for different
cell lines, which may respond differently to the treatment. This
data can be used to develop devices with optimized geometries and
operating parameters for specific (or specific subsets of) cell
types.
[0236] FIG. 38 depicts micrographs illustrating alternate device
structures. Bright field micrograph of preliminary work combining a
constriction and electrodes (scale bar 30 m). As illustrated in
FIG. 38, the device can be modified by coupling the rapid
deformation phenomenon with electroporation. Gold electrodes can be
incorporated on either side of the constriction by
photolithographic patterning and Au deposition to introduce a
localized electrical field into the channel. Subsequent experiments
can identify operating parameter values (electric field strength,
frequency, and operating speed) that demonstrate an improved
performance over current methods. By coupling two independent
poration mechanisms, one can exercise finer control over the system
and manipulate multiple parameters to optimize system performance
for each cell type. Additionally, electric fields can be used as a
driving force to deliver larger charged molecules, such as DNA,
that suffer from low diffusion rates.
[0237] A streamlined, disposable version of the system suitable for
use by potential collaborators is possible. Injection molding or
hot embossing of PMMA and polycarbonate can be used for
implementing a polymer-based version of the device. The subsequent
reduction in costs would enable these devices to be used as a
disposable tool hence improving sterility and ease-of-use. In
addition, by simplifying the tubing connections, mounting system
and pressure regulator setup, it is feasible to supply a
user-friendly system.
[0238] Protein-based reprogramming of fibroblasts into iPS cells
and optimization studies of the reprogramming parameter space is
possible. Previous studies have shown that iPSCs can be generated
by direct delivery of CPP-fused reprogramming factors from both
human and mouse tissues (FIGS. 36A-36F) and that these
protein-iPSCs can differentiate into functional cells (e.g.,
dopamine neurons) without abnormal phenotypes associated with viral
iPSCs. However, due to many factors including protein degradation
in culture, delivery inefficiencies, and degradation inside the
cells' endosomes, the reprogramming efficiencies by direct protein
delivery are too low (<0.01%) for any practical use. The
microfluidic devices described herein can significantly increase
the efficiency of protein-based reprogramming by allowing efficient
protein delivery directly to the cytoplasm thus avoiding the harsh
endosomal environment and the cumbersome endosomal escape process
usually encountered in free or encapsulated protein delivery
methods.
[0239] Generation of human iPSCs is facilitated by
microfluidic-based delivery. The device is used to deliver one or
more, e.g., 4 reprogramming proteins (c-Myc (protein, Genbank
Accession NP_002458.2; DNA, Genbank Accession NM_002467.4), Klf4
(protein, Genbank Accession AAH30811.1; DNA, Genbank Accession
NM_004235.4), Oct4 (protein, Genbank Accession ADW77326.1; DNA,
Genbank Accession HQ122675.1), and Sox2 (protein, Genbank Accession
NP_003097.1; DNA, Genbank Accession NM_003106.3);) to embryonic
human fibroblasts (HFs). In addition to Yamanaka four factor (MKOS
being c-Myc-K1f4-Oct4-Sox2), several additional factors (e.g.,
Lin28 (protein, Genbank Accession AAH28566.1; DNA, Genbank
Accession NM_024674.4) and Nanog (protein, Genbank Accession
AAP49529.1; DNA, Genbank Accession NM_024865.2), Esrrb (protein,
Genbank Accession AAI31518.1; DNA, Genbank Accession NM_004452.3),
Glis1 (protein, Genbank Accession NP_671726.2; DNA, Genbank
Accession NM_147193.2), and PRDM14 (protein, Genbank Accession
NP_078780.1; DNA, Genbank Accession NM_024504.3)) have been
identified to enhance reprogramming efficiency. It has been fully
established that mammalian expression and purification of 6 factors
(MKOS+Lin28 and Nanog; MKOSLN) and established the bioactivity of
each factor using reporter assays. These factors can be expressed
either in E. coli or in mammalian cells. Since E. coli-expressed
proteins lack post-translational modifications such as
phosphorylation, acetylation and ubiquitination, purified proteins
can be used following expression in mammalian cells (HEK293 and
CHO). E. coli-expressed proteins (commercially available from
Stemgent, Cambridge, Mass.) can be used for comparison. First,
FLAG-tagged reprogramming factors expressed in HEK293 cells by
transfection can be resuspened in a NP40 cell lysis buffer
containing 50 mM Tris-HCl, pH 7.4, 250 mM NaCl, 5 mM EDTA, 1%
NP-40, and protease inhibitors. Following centrifugation, collected
soluble fraction can be added with equilibrated anti-FLAG M2
agarose affinity gel. After washing with PBS twice, retained
FLAG-tagged proteins can be eluted by adding 0.1 mg/ml of FLAG
peptide (Sigma). Suspended solution of human fibroblasts and 4 or 6
purified proteins can be applied to the device and treated cells
can be plated on to 0.1% gelatin coated plates with conditioned
hESC media for 1, 2, or 3 days before the next delivery cycle.
After repeating protein delivery cycles (6-16) with the
microfluidic device, treated cells will be plated on mitomycin C
treated mouse embryonic fibroblast (MEF) and grown for 3-4 weeks
with regular hESC media. IPSC colonies can become visible within 3
weeks after seeding on the MEF. The efficiency of iPSC generation
can be compared to that by protein delivery using our original
CPP-fused recombinant proteins. These iPSC candidates can be
thoroughly examined for all criteria of authentic iPSCs, including
molecular and cellular properties as well as in vitro and in vivo
pluripotency, as previously described. Using 4 or 6 factors will
generate iPSC lines with much improved efficiency. It is also
possible to further express additional factors such as Esrrb,
Glis1, and PRDM14 and use these factors in reprogramming
experiments.
[0240] Reprogramming proteins and mRNAs and/or microRNAs can be
delivered in combination. The microfluidic device can be used to
deliver not only proteins but also any other macromolecules. To
take advantage of this unique property for optimal non-genome
integrating reprogramming, it is possible to combine the use of
reprogramming factors and mRNAs and/or microRNAs. In particular, it
is of great interest that iPSC lines can be successfully generated
by using only microRNAs. Indeed, lipofectamine-based transfection
of microRNAs can generate iPSC-like colonies. Since microRNAs
likely induce reprogramming in a different mechanism than the
reprogramming factors, appropriate combination of both
reprogramming proteins and microRNAs via the microfluidic device
can further enhance the reprogramming efficiency. The combined
delivery of proteins and mRNAs can significantly facilitate the
reprogramming efficiencies. Thus, an optimal combined treatment of
proteins, mRNAs, and/or microRNAs can be delivered using the
microfluidic device. Although microRNA/mRNA may not offer the
equivalent level of control as proteins, the device capacity for
high-throughput optimization studies still provides significant
gains in efficiency relative to previousl approaches.
[0241] The unique features of the microfluidic device can enable
delivery of various quantified amounts of each factor, in a
controlled, repeatable manner. Optimization of the current subject
matter facilitates the development of a reliable, high efficiency
tool for protein delivery to HFs. It is possible to elucidate the
optimal delivery quantities and frequencies of each reprogramming
factor. Unlike mRNA, plasmid or viral methods, the system does not
rely on the stochastic nature of gene expression and/or translation
to determine the effective intracellular concentration of
transcription factors. Thus, the device's ability to deliver
protein directly to the cytosol places it in a unique position to
exercise accurate control over the intracellular environment. A
series of delivery schedules are possible that vary the treatment
frequency (once every 1, 2, or 3 days) and protein concentration of
each of the four factors independently. In particular, based on
several reports indicating that higher levels of Oct4 is critical
for efficient reprogramming, it is possible to test the effect of
using different concentrations of Oct4 while keeping other factors'
concentrations the same. Different concentrations of c-Myc may be
evaluated for a given cell type because it has been found that its
high levels generate mostly transformed colonies instead of iPSCs
in some situations. Furthermore, it is possible to test the effect
of more frequent treatment of c-Myc due to its extremely short
half-life (.about.30 min) with appropriate concentration.
[0242] Optimization of temporal treatment of reprogramming factors
is facilitated using the described methods. Each factor has a
functional role and participates in the reprogramming process. At
least one and in some cases, combinations of factors are necessary
to achieve the desired reprogramming result. For instance, c-Myc is
known to suppress the expression of differentiation genes. In
addition, Klf4 is known to repress the microRNA let-7, which is
related to differentiation pathways and inhibition of pluripotency.
Thus, temporally regulated reprogramming can be possible by
treating c-Myc and/or Klf4 for initial period, based on a
suboptimal condition. In addition, although Nanog is not required
for iPSC generation, it is known to be crucial for final
establishment and maintenance of pluripotency. Thus, the effect of
adding Nanog at the later stage of reprogramming process can be
tested. Furthermore, the sequential treatment of microRNAs and
proteins can be tested and compared the reprogramming efficiency to
those by each treatment or simultaneous treatment. This temporally
regulated reprogramming is feasible due to a unique feature of the
microfluidic device and may be important to further optimize the
protein reprogramming. Reprogramming efficiency can be calculated
by dividing the number of colonies at day 28 by the number of
treated HF cells. Once completed, regression analysis can be used
to deduce the relative importance of each reprogramming factor, its
optimal concentration, optimal delivery frequency/timing and, as a
result, the optimal protocol to generate iPSCs. The ability to
control the amount and timing of protein delivered into each cell
can shed light on the functional significance of each factor in the
reprogramming process, thereby further enhancing the understanding
of cell reprogramming process and pluripotency establishment. In
addition, the results of this work can be used to further improve
device design to meet the specific demands of reprogramming and
enable the eventual development of clinically applicable
versions.
[0243] Using the optimized protein reprogramming procedure as
described above, the protocol can be generally applied to
patient-specific adult human fibroblast cells. Since the efficient
differentiation of ESCs and iPSCs into functional dopamine neurons
and the effects of transplantation has been studied, it is possible
to generate iPSCs or iPSC lines from human fibroblasts derived from
Parkinson's patients. Once iPSC lines are generated and
characterized, they are induced to differentiate into dopamine
neurons and characterize their cellular, molecular, physiological,
and electrophysiological properties. The dopamine neurons are are
tested for in vivo functionality following transplantation into
animal models of Parkinson's disease such as the genetic PD model,
aphakia mice.
[0244] Microfluidic-based protein delivery can be used for direct
cell conversion, e.g., the direct conversion of fibroblasts into
other cell types such as functional neurons, hepatocytes, and blood
cells. In the past, the manipulations used viral expression of key
transcription factors, causing significant chromosomal disruptions
and gene mutations, thereby highlighting the need to develop
non-viral, genome non-integrating conversion methods such as direct
protein delivery using the methods described here. Thus, the device
can be used for microfluidic-based protein delivery for direct cell
conversion. Since mammalian expression of certain transcription
factors are sometimes challenging, it may be more feasible to test
the conversion by one or two protein factors. However, it is
possible to convert fibroblasts to another cell fate using a single
factor e.g., Oct4 or Sox2 to generate blood or neural precursors
respectively. These proteins are readily available in a purified
form, for cell conversions using the microfluidic-based protein
delivery.
[0245] FIG. 44 is a bar graph illustrating GFP knockdown in HESCs
as measured by GFP intensity 48 hours after treatment with active
siRNA sequences and scrambled controls using the microfluidic
device and related methods. FIGS. 45A and 45B are two plots
illustrating the dye intensity and viability of human embryotic
stems cells after delivery of a 3 kDa blue dye.
[0246] Other embodiments are within the scope and spirit of the
invention. For example, due to the nature of software, functions
described above can be implemented using software, hardware,
firmware, hardwiring, or combinations of any of these. Features
implementing functions may also be physically located at various
positions, including being distributed such that portions of
functions are implemented at different physical locations.
[0247] It is noted that one or more references are incorporated
herein. To the extent that any of the incorporated material is
inconsistent with the present disclosure, the present disclosure
shall control. Furthermore, to the extent necessary, material
incorporated by reference herein should be disregarded if necessary
to preserve the validity of the claims.
[0248] Further, while the description above refers to the
invention, the description may include more than one invention.
* * * * *