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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5588_Библиотеки_им_академика_М_И_Перельмана
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166 Xuan Mu and Yu Shrike Zhang
death [101]. Extravasation is a dynamic process of tumor cells exiting from the lumen of
blood vessels to a new tissue; it is also a promising target for cancer immunotherapy [102].
The in vitro microfluidics-based vascular networks, as discussed in the previous section,
have shown great promise for exploiting the mechanism of tumor cell extravasation
[103–105]. In particular, this type of microfluidic device leverages the stable and
perfusable 3D microvascular networks within well-controlled microchannels, which
allows the manipulation of soluble cues, the recapitulation of a more physiological configuration, and real-time high-resolution imaging. Of note, these technical advantages of
microfluidics-based approaches are beneficial for the investigation of tumor cell extravasation and largely inaccessible to single layer- or membrane-based models [106–108] and
intravital microscopy-based animal models [109].
One pioneer work has adopted the on-chip microvascular network to investigate the
extravasation of breast tumor cells, MDA-MB-231 (Fig. 4A–C) [110]. The MDA-MB231 cells are perfused into the 3D microvascular network from lateral channels under low
shear stresses from 0.012 to 0.48 Pa. Of note, the on-chip microvascular network has
openings to allow the perfusion of cells and media. After 4-h inoculation, some
MDA-MB-231 cells are found to migrate out of the vascular lumen; others are crossing
the endothelial barrier and show protrusions contacting the matrix. The extravasation of
MDA-MB-231 starts from 30 min after perfusion and is almost completed within around
24 h. The reported time scale of extravasation is comparable to in vivo observations [111]
and other in vitro models [112]. Moreover, the configuration of microchannels enables
the high-resolution time-lapse imaging of the extravasation by standard confocal microscopy. The imaging results show that the extravasated cells can create a small opening
(approximately 1 μm) and to expand it (approximately 8 μm) on the endothelial barrier
without disrupting endothelial cell-cell junctions or compromising barrier functions.
The larger opening is sufficient to allow the nuclear transmigration and protrusions of
the MDA-MB-231 cells. The extravasation rate is associated with the inflammation of
the endothelial barrier and the types of cancer cells. Tumor necrosis factor-α (TNF-α,
2ngmL
1
)-induced inflammation increases the permeability of the endothelial barrier
by 2.1 folds, which leads to ruptures and thus facilitates the extravasation by 2.3 folds.
The extravasation rate of cancer cells also relates to the corresponding metastatic potential. For example, highly invasive HT-1080 fibrosarcoma cells exhibit an almost two-fold
increase in extravasation rate in comparison to MDA-MB-231 cells.
The chip-based microvascular network and microenvironment can be designed to be
organotypic [113]. Human bone marrow-derived mesenchymal stem cells (MSCs) and
osteoblast-differentiated cells (OBs) are added into the 3D gel to mimic the microenvironment of the bone marrow, which is characterized by the expression of bone-specific
markers, such as osteocalcin and bone-specific isoform of alkaline phosphatase.
A standard myoblast cell line, C2C12, is added to mimic the muscle microenvironment.
A specific clone of the metastatic breast cancer cells (BOKL-MDA-MB-231) is perfused
Fig. 4 (A) Schematics and a photo of the microfluidic device with the microvascular network. (B) Confocal fluorescence images of MDA-MB-231
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cells in the middle of extravasation. The protrusions penetrate the HUVEC barrier and contacting the matrix. (C) The microvascular network is free
of leaks, as evidenced by the perfusion of 70-kDa dextran. (D) Schematic of the microfluidic device and the configuration of cell seeding. The
microvascular network is formed within the center channel with the assistance of fibroblasts. The monocytes and tumor cells are perfused into
the microvascular network. The tumor cells may migrate outside the microvascular network and form colonies in the matrix. (E) The addition of
circulatory monocytes reduces the extravasation rate of tumor cells after 5 h of incubation. (F) Fluorescence confocal image of both tumor cells
(red) and monocytes (white) within the microvascular network (green). (A–C: Reproduced with permission from Chen MB, Whisler JA, Jeon JS, Kamm
RD. Mechanisms of tumor cell extravasation in an in vitro microvascular network platform. Integr Biol 2013;5(10):1262–1271. Copyright 2013 Oxford
University Press. D–F: Reproduced with permission from Boussommier-Calleja A, Atiyas Y, Haase K, Headley M, Lewis C, Kamm R. The effects of monocytes on tumor cell extravasation in a 3D vascularized microfluidic model. Biomaterials 2019;198:180–193. Copyright 2019 Elsevier.)
168 Xuan Mu and Yu Shrike Zhang
into the microvascular network and the cells are found to roll on and adhered to the
endothelium, spread, extend filopodia through intercellular junctions, and finally, invade
into the ECM to complete the process of extravasation. Notably, the extravasation rate of
MDA-MB-231 cells shows a strong dependence on the organotypic microenvironments.
The extravasation rate in the bone marrow-mimicking microenvironment is 56.5%
4.8% that is much higher than that of the muscle-mimicking microenvironment
(8.2% 2.3%). The difference of cancer cell extravasation between bone marrowmimicking and muscle-mimicking microenvironments is associated with the factor
secreted by organ-specific cells. It is known that the skeletal muscle microenvironment
reduces cancer cell tumorigenicity via adenosine and A
adenosine receptor (A3AR)
3
[114,115]. The addition of adenosine to the bone marrow-mimicking microenviron-
ment reduces the extravasation rate from 56.5% 4.8% to 12.7% 2.8%. In contrast,
the addition of an A
AR-antagonist, PSB-10, to the muscle-mimicking microenviron-
3
ment increases the extravasation rate from 8.2% 2.3% to 32.4% 7.7%. These results
verify the critical role of adenosine in cancer cell extravasation and highlight the utility of
microfluidic-based microvascular networks for investigating the mechanisms of extravasation and for screening drug candidates.
The microfluidics-based microvascular network has been exploited to examine the
effects of other factors on cancer cell extravasation, such as monocytes [116] and luminal
fluid flows [117]. Intravascular and mobile monocytes can reduce the extravasation of
MDA-MB-231 by 42% within 5 h after perfusion (Fig. 4E and F) [116]. In contrast,
the monocytes in the macrophage-like morphology in direct contact with cancer cells
cannot reduce the extravasation rate. This result suggests that the reduction of the extravasation rate by intravascular monocytes results from the paracrine communications
between the two types of cells. In particular, monocytes secrete a considerable amount
of tissue inhibitors of matrix metalloproteinases (TIMP)-1 and TIMP-2, the overexpressing of which has been proposed to reduce cancer cell extravasation by preventing
cancer cells from breaching the basement membrane [118]. Monocytes in the patrolling
state have been shown to reduce metastasis by promoting the recruitment and activation
of natural killer cells and the scavenge of tumor material [119] . Thus, the intravascular
monocytes to reduce cancer extravasation in a paracrine manner may represent a new
strategy of cancer immunotherapy [116]. Besides, the presence of luminal flow promotes
the extravasation rate of tumor cells [117]. The increase of transendothelial fluid increases
the migration speed of cells across the endothelium and into the matrix [117].
3.3 Lymphocyte trafficking
Lymphocyte trafficking, into and within lymphoid and peripheral tissues, plays important
roles in immune cell development, immune responses, immunosurveillance, as well as
cancer immunotherapy [120,121]. In particular, the infiltration of T cells from peripheral
tissues into the solid tumor usually represents a critical hurdle responsible for ineffective
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immunotherapeutic outcomes [122]. The migration and positioning of lymphocytes
have been known to be regulated by chemokines, i.e., chemotaxis [120]. For example,
the infiltration of immune cells into melanoma tumors is often mediated by the tumorsecreted chemoattractant, such as CdC Motif Chemokine Ligand 2 (CCL2), CdXdC
motif ligand 9 (CXCL9), CXCL10, and CXCL11 [123]. In comparison to conventional
methods, such as transwell assays, microfluidics has demonstrated outstanding capabilities
of precisely manipulating the patterns and gradients of chemokines, automated analyses,
and high(er)-throughput, which are promising for the investigation of the chemotaxis of
immune cells [124,125].
A microfluidic chip with the classic pyramidal branched array of microchannels has
been developed for the neutrophil chemotaxis in the gradient of IL-8 (Fig. 5A and B)
[126]. The IL-8 induced chemotaxis of neutrophil is mediated by a nonreceptor tyrosine
kinase, Janus kinase 3 (JAK3) [127]. The microfluidic chip has two inlets, one for Hank’s
balanced salt solution (HBSS) and the other for IL-8 solution (50 ng ml
1
). The streams
from the two inlets are divided, combined, and mixed in a repeated manner when the
flow passes the pyramidal network of branched channels (50 -μm-wide) under the laminar flow condition with a low Re number. The stream from each channel with a distinct
concentration of IL-8 eventually converges into the main channel (500 -μm-wide) to
give rise to a linear gradient of IL-8 that is perpendicular to the flow direction and
maintained several millimeters long within the main channel. The gradient is stable in
a spatial-temporal manner yet requires continuous flow and imposes low shear stress
(<0.1 dynes cm
2
) on cells. Other alternative microfluidic devices enable the generation
of stable gradients without continuous flow, such as the use of a hydrogel over the microfluidic channels [128].
By using multiple microfluidic chips, more complex patterns of the IL-8 gradient can
be realized. Two parallelly placed chips result in a hill-type gradient, where the highest
concentration of IL-8 (50 ng mL
1
) stays in the middle of the main channel, and minimal
IL-8 is at the boundary of the channel. In another configuration, two microfluidic chips
give rise to a cliff-type gradient, where the concentration of IL-8 increases from 0 to
50 ng mL
1
, decreases to 0 ng mL1, and increases again to 50 mg mL1. Also, the
direction and magnitude of the gradient can be readily tuned by changing the input liquid
and flow rate without altering the design of the microfluidic chip. The flexibility of chemokine gradients in the microfluidic chip highlights the promise to investigate the chemotaxis of immune cells, in contrast to other in vitro techniques, including Boyden
chamber [129] and Dunn chamber [130].
The microfluidic device reveals new insights into the migration behavior of neutrophils in the complex cliff- and hill-like gradients of IL-8 (Fig. 5B). In the linear-gradient,
neutrophils demonstrate unidirectional migration toward the high concentration of IL-8.
In the hill-like gradient, neutrophils migrate toward the maximal concentration of IL-8
169Tumor-on-a-chip devices for cancer immunotherapy
170 Xuan Mu and Yu Shrike Zhang
Fig. 5 (A) Schematic of the microfluidic chip with a pyramidal branched array for generating an
in-channel gradient of chemoattractant and for the investigation of neutrophil chemotaxis. (B) One
linear and two nonlinear gradients can be generated by using one or multiple microfluidic chips.
(C) Image of the microfluidic chip for the high-throughput screening and isolation of population-based
chemotactic neutrophils. Two dye solutions (blue and violet) are perfused into the channel to demonstrate the transverse gradient. (D) Typical images of neutrophils in the five outlet chambers.
(E) Quantitative analysis of the neutrophil distribution in the five outlet chambers in the absence
(top) and the presence (bottom) of a gradient of 100-ng mL
1
CXCL2. (F) Illustration, fluorescence
images, and line profiles of the competing gradients of two components. (G) Track lines of T cells’
migration under the competing gradient of CCL19 at the right and CXCL12 at the left. Darker tracks
indicate the direction that is toward the right of the channel, and lighter tracks toward the left. (A and B:
Reproduced with permission from Jeon NL, Baskaran H, Dertinger SKW, Whitesides GM, Van de Water L,
Toner M. Neutrophil chemotaxis in linear and complex gradients of interleukin-8 formed in a microfabricated device. Nat Biotechnol 2002;20(8):826–830. Copyright 2002 Springer Nature. C–E: Reproduced
with permission from Grigolato F, Egholm C, Impellizzieri D, Arosio P, Boyman O. Establishment of a scalable microfluidic assay for characterization of population-based neutrophil chemotaxis. Allergy 2020;75
(6):1382–1393. Copyright 2020 John Wiley and Sons. F and G: Reproduced with permission from Lin F,
Butcher EC. T cell chemotaxis in a simple microfluidic device. Lab Chip 2006;6(11):1462–1469. Copyright
2006 Royal Society of Chemistry.)
in the middle of the channel; the neutrophils often pass beyond the location of the max-
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imal concentration but reverse their directions of migration back to the maximal concentration. In contrast, neutrophils rarely overshoot the maximal concentration in the cliff-like
gradient that has a sharp transition from 50 to 0 ng mL
1
. These data at the single-cell resolution indicate that the neutrophils are adaptive in the nonlinear gradients of the chemoattractant, especially the steep gradient, highlighting the promise of microfluidics to
manipulate nonlinear gradients for studying lymphocyte chemotaxis [131].
In addition to the analyses at the single-cell resolution [126], another microfluidic
device enables the high-throughput screening and isolation of population-based chemotactic neutrophils (Fig. 5C–E) [132]. The neutrophils demonstrate a higher lateral migration speed in the presence of the gradient of CXCL2 (11.9 8 μm min
the absence of the gradient (0.3 3.2 μm min
1
). The lateral migration of neutrophils
1
) than that in
enables the collection of cells in different outlets with distinct lateral positions. This result
of CXCL2-induced chemotaxis of neutrophils is confirmed by blocking the CXCL2 signaling pathway. The neutrophils without CXCR2 and the incubation of an inhibitor,
IL-4, lead to random migration that is nondifferentiable from controls.
Besides neutrophils, the chemotaxis of human peripheral blood T cells is investigated
in microfluidic devices with competing gradients of two chemokines ( Fig. 5F and G)
[133]. The chemotactic cytokines, CCL19 and CXCL12, are patterned in opposite
directions within the channel. Both the ratio of chemotactic cells and the migration speed
are higher in the gradient of single chemokines than that of the competing gradients. This
result shows that immune cells can respond to multiple chemokines. Also, CCL19 results
in a higher number of chemotactic cells and a faster migration speed than CXCL12,
implying specific effects of chemokines on the signaling pathway of immune cells
(Fig. 5G).
171Tumor-on-a-chip devices for cancer immunotherapy
4. Predicating therapeutic efficacy
Precision or personalized medicine requires the prediction of the therapeutic
response of individual patients based on not only genomics but also functional features
[34,134]. Microfluidics offers the capability to engineer tumor spheroids and exploit
native tissue samples, which has shown tremendous advantages in the prediction of
responses of the individual patients [135].
In contrast to tumor spheroids consisting of only tumor cells, bioengineered patient-
or murine-derived organotypic tumor spheroids (MDOTS/PDOTS) contain autologous immune cells and have been used in microfluidic chips to examine the ex vivo
response and efficacy of immune checkpoint blockade (Fig. 6A–C) [136,137]. The
immune cells, infiltrated inside the tumor, are important components of the tumor
microenvironment; the inclusion of these immune cells in the tumor spheroids is
Fig. 6 (A) Fluorescence image of PDOTS from high-grade serous carcinoma (HGSC) of the ovary. Tumor and T cells are labeled by EpCAM (violet)
and CD8 (red), respectively. (B) MDOTS/PDOTS in the diameter from 40 to 100 μm are pelleted and suspended in collagen, followed by injection
into microfluidic channels for ex vivo culture of 5–9 days. A range of methods can be used to characterize and monitor the MDOTS/PDOTS in
response to the immunotherapy. (C) The dual immune checkpoint blockade (αPD-1 + αCTLA-4) leads to the highest percentage of dead tumor
cells than that of the control and the single immune checkpoint blockade. (D) Schematic and images of the microfluidic chip for multiplexed drug
testing of tumor slice. The multiplex and parallel microfluidic channels or drug lines are highlighted in three dyes, including cell tracker red, cell
tracker green, and hoechst (blue). (E) Apoptotic cell death (CC3 staining) in response to different drugs. The on-device results are consistent with
the off-device ones. (A–C: Reproduced with permission from Aref AR, Campisi M, Ivanova E, Portell A, Larios D, Piel BP, Mathur N, Zhou C, Coakley RV,
Bartels A. 3D microfluidic ex vivo culture of organotypic tumor spheroids to model immune checkpoint blockade. Lab Chip 2018;18(20):3129–3143.
Creative Commons license. D and E: Reproduced with permission from Horowitz LF, Rodriguez AD, Dereli-Korkut Z, Lin R, Castro K, Mikheev A, Monnat
RJ, Folch A, Rostomily RC. Multiplexed drug testing of tumor slices using a microfluidic platform. NPJ Precis Oncol 2020;4(1):1–15. Creative Commons
license.)
expected to recapitulate sensitivity and resistance to immune checkpoint blockade ex
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vivo. The presence of immune cells within the MDOTS/PDOTS is verified by the
immunostaining of specific markers, such as CD 8 (Fig. 8A) [138]. The MDOTS/
PDOTS are selected with the diameter in the range from 40 to 100 μm, suspended in
collagen solution, followed by perfusion into microfluidic channels for ex vivo cultures
and characterizations [138]. The gel-laden microchannel in each chip is flanked by
two lateral channels for perfusion of culture media, separated by micro-posts. In comparison to other in vitro models, such as patient-derived xenografts, the adoption of
microfluidic chips is promising for scaling up, manipulating soluble cues, enhancing
experimental reproducibility, and enabling real-time imaging and rapid screening
[38,47,138]. The death of tumor cells is a primary readout for PD-1 blockade, which
is observed in the microfluidic MDOT model but not in the 2D model. Furthermore,
in a PDOTS model of small intestinal neuroendocrine tumor, a dual immune checkpoint
blockade with both αPD-1 and αCTLA-4 leads to a higher percentage of cancer cell
death, in comparison to single-agent PD-1 blockade or CTLA-4 blockade. Highsensitivity bead-based immunoassay is used to examine the secretion of cytokines and
growth factors by the MDOTS/PDOTS over the 9-day culture in the chip. The cytokine profile is associated with immunotherapy strategies, for example, dual or single
immune checkpoint blockade. The compositional change of immune cells is also examined by RNA sequencing (RNA-seq) and a computational method, cell-type identification by estimating relative subsets of RNA transcripts (CIBERSORT) [138].In
particular, the dual immune checkpoint blockade leads to a much higher portion of
CD8 T cells and monocytes in comparison to that of the single-agent blockade,
suggesting the higher efficacy of the dual immune checkpoint blockade. The capability
to test the efficacy of immunotherapy strategies is particularly beneficial to enhance longlasting disease control for vast patients [134].
Functional drug screening has been exploited with a multiplexed microfluidic device
and patient-derived intact tumor slices (Fig. 6D and E) [139,140]. This microfluidic
device consists of 40 parallel channels, separated by 500 μm, and is made in a thermoplastic poly(methyl methacrylate) (PMMA). The adoption of PMMA eliminates the
common issues of PDMS, for example, the bulk absorption of drug molecules
[93,141,142]. Xenograft-derived or patient-derived tumor slices (approximately 250-μ
m-thick) are cultured on a polytetrafluoroethylene (PTFE) porous membrane that is then
transferred onto the 40 microfluidic channels. The use of tumor slices is due to the
retaining of the original tumor’s heterogeneous content and 3D structures, including
ECM, stromal and immune cells, and biochemical cues [143]. The tumor slice can be
further processed into “cuboids” which are hydrodynamically trapped in a microfluidic
chip to form a high throughput array [140].
The 40 parallel channels can deliver multiple drugs or concentrations to distinct loca-
tions of a single tumor slice (Fig. 6D and E) [139]. The delivery of drugs is primarily
173Tumor-on-a-chip devices for cancer immunotherapy
174 Xuan Mu and Yu Shrike Zhang
through diffusion, which leads to a width of 55 μm after 48-h culture. The spatially controlled delivery of drugs not only takes full advantage of the precious tumor slice but also
improves the efficiency and reproducibility of drug screening. As a result, the efficacy of
drugs in inducing apoptotic cell death, stained by cleaved-caspase 3 (CC3), can be examined (Fig. 6E). In addition, xenograft tumor slices derived from the flank and intracranial
(IC) tumors show different drug responses, indicating the importance of the tumor
microenvironment. Although only small molecules are tested, the tumor slice-based
microfluidic platform seems equally useful for screening drug candidates and strategies
of cancer immunotherapy.
5. Production of therapeutic cells
Besides mimicking the tumor microenvironment, microfluidics has been extensively exploited for producing therapeutic cells and materials for cancer immunotherapy.
In comparison to conventional approaches, microfluidic ones are often characterized by
higher efficiency, shorter operation time, and versatile cellular functions. The production
of therapeutic cells has been an essential topic related to tumor-on-a-chip.
5.1 T cells
The adaptive T-cell immunotherapy [144], including both chimeric antigen receptors
(CARs) and T-cell receptors (TCR), relies on the delivery of plasmids and gene-editing
tools into the T cells to boost the recognition of cancer antigens, thus enabling specific
attacks against cancer cells. However, the delivery efficiency, depending on the delivered
materials and cell types, often has been a challenge. For example, the ablation efficiency of
Cas9 and single-guide RNAs (sgRNAs) is only 1%–5% of the target protein expression in
human CD4 + T cells [145]. Furthermore, conventional delivery approaches are limited
in several aspects. First, viral plasmid-mediated delivery [146,147] is often confined to
RNAs and DNAs and may lead to uncontrolled chromosomal integration to the host
genome, thus eliciting adverse immune responses [148,149]. Second, as the cell membrane is largely impermeable to macromolecules, some auxiliary materials, such as mesoporous silica nanoparticles [150], lipid nanoparticles [151,152], and cell-penetrating
peptides [153], have to be used to assist the delivery via endocytosis. Nevertheless, these
materials introduce exogenous materials and still suffer from a low transfection efficiency
for primary T cells. Third, electroporation [154] and sonoporation [155] are usually of
high efficiency for delivery yet tending to induce irreversible cellular damage and reduce
cellular viability [156,157]. Fourth, microinjection is attractive for delivering materials to
single cells for producing transgenic organisms. However, this technique suffers from
sophisticated instruments and low throughput, compromising the potential for therapeutic applications that require a large number of cells [158]. Thus, new delivery methods
with high efficiency and minimal cellular damage would be highly desired to produce
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therapeutic T cells for cancer immunotherapy.
To this end, a series of microfluidic approaches are promising, which exploit the mild
deformation of cell membranes to achieve cytosol deliver gene-editing tools (CRISPR/
Cas9) into immune cells [159–161]. In particular, the strategy to enable mild membrane
deformation is to let cells rapidly pass a geometric constraint (for example, with 40 μmin
length and 6 μm in width, latter smaller than the size of cells) [159]. Due to the geometric
constraint, the cells are mechanically deformed and, most likely, elongated, leading to
transient and revisable membrane disruption. The disruption of cell membranes allows
the delivery of surrounding materials into the cytosol primarily via passive diffusion.
The throughput of membrane-deformation-based delivery is up to 20,000 cells per
second, allowing the delivery to around one million cells at one time. Also, the membrane deformation-based delivery eliminates the need to use a range of conditions unfavorable for cellular viability, including electric fields, exogenous materials, active
endocytosis, and chemical modifications.
The delivery efficiency and cell viability are associated with the flow velocity and the
geometry of the constriction, including the length, width, and number of repeated units
[159]. The increase of flow velocity from 0 to 600 mm s
1
leads to the decreased viability
of HeLa cells from around 97% to 90% and the increased delivery efficiency of cascade
blue-conjugated 3-kDa dextran from roughly 5% to 80%. Of note, even at the highest
flow rate (600 mm s
1
), the cell viability remains high (90%). The remained high cell
viability is ascribed to the optimized geometry of the physical constriction, which minimizes the magnitude and duration of the exerted forces on cells. Of note, for transfecting
primary dendritic cells, a difficult-to-transfect cell, the delivery efficiency of cascade blueconjugated 3-kDa dextran is around 40%, higher than other conventional methods [159].
Passive diffusion, different from active endocytosis, is proposed to be the primary
mechanism in the membrane deformation-based delivery [159]. Several pieces of experimental evidence have been used to support this argument. The cascade blue-conjugated
3-kDa dextran, after delivery into the cells, demonstrates a relatively uniform pattern, in
contrast to the punctate characteristics caused by endocytosis. The delivery is also completed roughly within 1 min, which is consistent with the membrane sealing that occurs
after a disruption (around 30 s) [162]. The diffusion-dominated delivery is subject to the
concentration gradient, according to Fick’s first law. Indeed, after flowing through the
geometric constriction, the dextran-laden cells show the decreased content of the dextran. This result is ascribed to the passive diffusion of dextran from cells to the surrounding environment due to the concentration gradient of dextran. Despite all these results,
passive diffusion may not be the sole mechanism involved in the delivery, as the delivery
efficiency is not affected by the temperature. While the diffusion coefficient relates to
temperature, according to Stokes–Einstein equation [163]; a lower temperature leads
to shorter diffusion distance and thus less efficiency of the diffusion-based delivery.
175Tumor-on-a-chip devices for cancer immunotherapy
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