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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5346_Библиотеки_им_академика_М_И_Перельмана
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74 Khushbu Bhatt et al.
disadvantages. The implantation of scaffolds often requires an invasive surgery and cannot
be achieved in surgically inaccessible and volume-sensitive locations [49]. Furthermore,
their prolonged presence may impair normal physiological tissue or organ function
[49,62]. These shortcomings led to the development of injectable polymeric scaffolds
for immunotherapeutic purposes. Compared to surgical procedures, injectable scaffolds
are less invasive and can be utilized in all areas where a needle can reach [49]. This can
prevent tissue damage and postoperative surgical wound infection [49]. Additionally,
minimal technical expertise is required for injections in comparison to surgical implantations [60]. However, a major disadvantage of injectable systems is that when injected as
a pregel solution, they must rapidly solidify upon injection, or when injected as a solid,
they must either momentary liquify or reversibly collapse to pass through small-bore
needles, thus restricting the type of polymers and components that can be incorporated
[49]. Various natural and synthetic polymers such as chitosan, HA, alginate, collagen, gel-
atin, PEG, PLGA, and PNIPAAm can be used to fabricate injectable scaffolds such as
conventional hydrogels and cryogels [49,63–65]. In situ-forming hydrogels are injected
in a liquid form and subsequently transition to a solid scaffold inside the body via chemical
or physical crosslinking mechanisms [66–68]. However, this approach carries certain limitations, such as the need to maintain appropriate gelation time and conditions, the
potential formation of gel scaffolds with poor mechanical properties, and the inability
to protect the encapsulated cargo in adverse biological environments [69]. Moreover,
there is potential for the precursor liquid solution to either leak into the surrounding tissue or be diluted within the body fluids, leading to gel formation with inadequate physical
properties [70–72]. To circumvent these limitations, solid preformed hydrogels such as
shear-thinning hydrogels have been developed using physical crosslinking; they liquefy
under shear-stress and, once injected, regain their solid gel-like properties [73,74]. However, these hydrogels often possess poor mechanical properties due to the weak nature of
physical crosslinking in addition to a quasi-nonporous structure (i.e., mesoporous) and
lack of geometrical shape memory [73]. To improve upon this, Bencherif et al. developed
injectable cryogel scaffolds possessing shape-memory properties, a macroporous and
interconnected network, and enhanced elastic mechanical properties, including a high
degree of compressibility allowing straightforward injection through hypodermic needles
[69,75–77]. Since these preformed scaffolds possess viscoelastic properties, they can easily
move and flow, and conform to any accessible space before establishing a durable implant
[78]. Most shape-memory scaffolds can be injected through large needles. However,
injecting them through finer needles remains challenging. Moreover, the use of large
needles (<18G) may lead to tissue damage and injury. Overall, both implantable and
injectable polymeric scaffolds have their distinct advantages which can be exploited
for desired applications accordingly. Furthermore, the ever-changing scope of research
is not just limited to implantable or injectable polymeric scaffolds for immunotherapy.
Chen et al. employed a sprayable polymeric system. Specifically, they used a
bioresponsive fibrin-based hydrogel that was formed in situ by spraying a thrombin and
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fibrinogen solution containing CaCO
nanoparticles loaded with CD47 antibody [79].
3
When introduced into an acidic TME, which mainly results from the high glycolytic rate
of tumor cells, the resulting nanocomposite gel system gradually released CD47 antibody
as a targeted therapy against cancer [79].
3. Polymeric scaffolds for DC priming and activation
DCs are professional APCs that function at the interface of the innate and adaptive
immune responses. DC-based vaccination strategies have been explored to induce
tumor-specific T cells that not only cause tumor regression but also generate a memory
response to avoid future relapse. Interestingly, adoptive DC transfer has been shown to
induce T-cell responses in vivo [80,81]. However, these approaches are hindered by
impaired cell survival and inadequate homing of DCs to the lymph nodes, resulting in
low efficacy [82–86]. Moreover, tumor-mediated inhibition induces the transition of
adoptively transferred DCs from immunostimulatory to immunosuppressive phenotypes,
leading to immune evasion [87]. Therefore, there is a need for alternative strategies for
DC transfer to circumvent tumor-induced immunosuppression and promote effective
lymphocyte recruitment and activation. Three-dimensional macroporous scaffolds can
act as delivery vehicles for adoptive DC transfer to a specific site. Moreover, they can
be functionalized with additional cues to promote DC activation, proliferation, and survival, along with their steady release into the local environment (Fig. 1). Hori et al. provided some initial evidence using an injectable alginate-based in situ-gelling hydrogel to
deliver activated antigen-loaded DCs [88–90]. In their approach, sacrificial calciumcrosslinked alginate microspheres (i.e., porogens) were mixed with alginate and DCs
to obtain a solution that polymerized immediately upon injection in the body, forming
a macroporous hydrogel. The authors reported that the delivered DCs recruited endogenous DCs and T cells at the injection site whereas a subset of injected DCs traveled to
the draining lymph node. This strategy facilitated the priming of naı¨ve T cells, which
subsequently trafficked into the hydrogel due to the local inflammatory microenvironment created by the scaffold [89]. In their subsequent study, the authors incorporated
IL-15 superagonist together with activated DCs into the self-gelling alginate scaffold,
which enhanced the recruitment of CD8+ T cells. Hydrogel-mediated delivery of peritumoral DCs and IL-15 superagonist was more efficacious than free DCs that were
directly injected into the tumor, thus highlighting the importance of biomaterials. Moreover, unlike bolus injections, the peritumoral injections of hydrogels achieved a 40-fold
higher local concentration of IL-15, with lower levels in the systemic circulation. Furthermore, this approach limited any potential adverse side effects while accumulating
IL-15 in the tumor milieu. Additionally, since the preparation of activated DCs is a
75Polymeric scaffolds for antitumor immune cell priming
76 Khushbu Bhatt et al.
Fig. 1 Schematic illustration of polymeric scaffold-based adoptive transfer of dendritic cells. Large
numbers of peripheral blood monocytes are collected from leukapheresis procedures from cancer
patients. These monocytes are differentiated into immature DCs ex vivo, using IL-4 and GM-CSF. Subsequently, ex vivo differentiated DCs are loaded into the polymeric scaffolds, which are incorporated
with DC growth factors, adjuvants, stimulatory cues such as tumor cell lysate or antigenic peptides, and
T-cell recruiting factors. This provides an artificial microenvironment that supports DC survival, activation, proliferation, and effector functions. Upon administration of the polymeric scaffold near the
tumor, CD8+ T lymphocytes are recruited into the scaffold. Activated antigen-loaded DCs interact with
recruited naïve CD8+ T cells and facilitate T-cell priming inside the scaffold. Furthermore, these exogenous DCs expand within the scaffold and are slowly deployed into the TME. A fraction of injected DCs
also travels to the nearby draining lymph nodes, thereby mounting effective local as well as systemic
antitumor immune responses.
labor-intensive and costly process, the authors loaded TLR-9 agonist CpG-ODN with
IL-15 superagonist in DC-free alginate hydrogels to improve the recruitment and activation of host DCs. The rejection of melanoma tumors with peritumoral alginate gels
carrying DCs + IL-15 superagonist was equivalent to a single dose of alginate gels loaded
with IL-15 superagonist + CpG and two doses of alginate gels loaded with IL-15 superagonist alone, thus excluding the need for exogenous DCs and simplifying this
approach [88]. More recently, Verma et al. utilized biodegradable and macroporous
fibrin-based scaffolds to deliver activated DCs in mice. Transferred DCs led to a significant tumor regression of primary and postsurgery residual tumors when compared to the
control group (i.e., transferred scaffold-free DCs). Furthermore, 50% of mice with complete remission survived a subsequent tumor challenge, suggesting long-lasting antitumor
immunity. Interestingly, DC-harboring scaffolds recruited host DCs as well as CD8+
T cells. The CD8+ T cells were found to interact with the exogenous DCs, indicating
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that T-cell priming may have occurred within the scaffold [91]. Taken together, these
studies have demonstrated that polymeric matrices can be used as immune cell depots
to effectively deliver DCs around tumors, thereby bypassing tumor-associated immunosuppression for an effective local and systemic antitumor immune response.
In contrast to ex vivo activation and adoptive transfer, DCs can be directly modified
in vivo, thus avoiding the need for expensive, time-consuming, and laborious ex vivo
manipulation of cells. Macroporous 3-D polymeric scaffolds are excellent tools for in situ
recruitment and modulation of immune cells due to their high surface area and spatiotemporal release of immunomodulatory factors. This strategy can generate effective
immune responses for cancer treatment. In their seminal work in 2002, Kumamoto
et al. designed an innovative approach to modulate endogenous DCs by engineering a
polymeric scaffold to release DC-recruiting chemokines and tumor cell lysate. They subcutaneously implanted PEVA-based rods encapsulating macrophage inflammatory protein (MIP)-3β. In a second step, the authors coimplanted the rods loaded with tumor cell
lysate, artificial antigens such as ovalbumin (OVA), or MHC-I restricted peptides.
Approximately 70% of the encapsulated MIP-3β was released in 48 h, leading to the infiltration of Langerhans cells, a type of skin DCs, into the polymeric matrix. Application of
haptens such as dinitrofluorobenzene over the implant site facilitated the maturation and
release of Langerhans cells from the epidermis to the local draining lymph nodes. Furthermore, the authors reported the efficacy of their in situ vaccine in EG7-OVA tumor,
fibrosarcoma, and carcinoma, both prophylactically and therapeutically [92]. In a more
advanced approach, Liu et al. engineered a minimally invasive and injectable
thermosensitive hydrogel fabricated from monomethoxy PEG-PLGA (mPEG-coPLGA) copolymer. In the first step, the hydrogel released GM-CSF in a sustained manner
to recruit and activate DCs and macrophages at the injection site. In the second step,
cancer antigens, encoded in viral and nonviral vectors, were delivered at the injection
site to stimulate recruited DCs and facilitate CTL responses. This strategy induced
antigen-specific T-cell responses and significantly improved the survival rate of mice
in a melanoma model, both prophylactically and therapeutically [93].
Compared with this two-step approach, Mooney and colleagues conceived a simplified one-step strategy by designing various implantable and injectable polymeric scaffolds
for the simultaneous delivery of a DC-recruiting factor and tumor antigens. Ali et al.
designed a macroporous and biodegradable PLGA-based implantable scaffold encapsulating GM-CSF to recruit DCs, CpG oligodinucleotides (CpG-ODN) to provide the
danger signal, and tumor cell lysate as a source of tumor antigens. Upon subcutaneous
surgical implantation, the PLGA scaffolds enabled sustained release of GM-CSF to
recruit DCs in a dose-dependent manner. Polyethyleneimine (PEI)-condensed CpGODN recruited plasmacytoid DCs and stimulated their activation and maturation,
77Polymeric scaffolds for antitumor immune cell priming
78 Khushbu Bhatt et al.
thereby promoting their subsequent homing to the draining lymph nodes. Robust
antigen-specific CD8+ T-cell responses were observed in a murine melanoma model
which led to 90% efficacy in a prophylactic setting and 50% efficacy in a therapeutic setting. Notably, these findings highlight the significance of providing a porous scaffold to
recruit and train DCs, since bolus administration of the same immunomodulators resulted
in lower efficacy [82,94]. Furthermore, the authors established the efficacy of their
PLGA-based vaccine in a rat intracranial glioma model and a murine lung carcinoma
model, highlighting the broad applicability of this approach [95,96]. In their subsequent
studies, Ali et al. demonstrated the versatility of these polymeric matrices to deliver
different cytokines and chemokines, such as Flt3L and CCL20, and various immune
adjuvants such as poly I:C and MPLA [96,97]. Additionally, they combined their
PLGA-based vaccine with intraperitoneal administration of immune checkpoint antibodies such as anti-CTLA-4 and anti-PD-1 to reinforce cytotoxic T-cell activity and
promote tumor rejection in an aggressive murine B16F10 melanoma model [98].Ina
different approach, Choi et al. fabricated mesoporous silica-based 3-D scaffolds with
strong mechanical properties and a bimodal porous structure. Mechanically enhanced
GM-CSF-loaded scaffolds recruited a significantly higher number of myeloid cells and
DCs as compared with mechanically weaker scaffolds [99]. More recently, Sinha et al.
designed a 3-D alginate-graphene scaffold as a vaccine delivery platform for long-term
activation of DCs. Scaffolds integrated with reduced graphene oxide offered a high surface area and a hydrophobic surface which facilitated increased loading and a slow and
sustained release of GM-CSF and model antigen OVA. These characteristics of the scaffold aided the recruitment and activation of DCs for up to 30 days after implantation due
to the prolonged availability of encapsulated immunomodulators. Furthermore, the anticancer vaccine induced effector CD4+ and CD8+ T-cell responses as well as memory
T cells which suppressed B16-OVA tumors in a prophylactic setting [100]. Overall,
implantable macroporous polymeric scaffolds have a unique capability to persist longterm at the implanted location while continuously releasing immune modulators, thereby
promoting cellular trafficking into the scaffold. Implantable scaffolds can be engineered to
fine-tune the release kinetics of immunomodulatory factors to further control immune
cell recruitment and activation. Such robust and durable depots have produced striking
results in various preclinical tumor models, and they are currently being evaluated for
their potential clinical translation [49]. Despite their ability to stimulate potent anticancer
immune responses, implantable polymeric scaffolds are limited in their widespread application as they require invasive surgical interventions when introduced into the body.
Administration of injectable polymeric matrices is more straightforward and safer
when compared to invasive surgical procedures needed for implantable scaffolds. Injectable scaffolds can be made from in situ-forming or deformable preformed hydrogels. For
instance, Singh et al. reported an in situ-forming biodegradable hydrogel-based delivery
system fabricated using dextran vinylsulfone and tetra-thiolated polyethylene glycol. The
scaffold encapsulated DC-attracting chemokines (e.g., CCL20) along with PLGA micro-
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particles which were loaded with short interfering RNA (siRNA) targeting IL-10 and
plasmid DNA antigen. Upon injection, the hydrogel-based formulation formed a porous
immune-priming niche, that promoted trafficking and programming of DCs, and significantly improved CTL responses over a naked DNA vaccine. Such a multicomponent,
injectable system provided a significant survival benefit of up to 2-fold increase in an A20
B-cell lymphoma model [101,102]. Mesoporous silica rods (MSRs) are another type of
injectable polymeric scaffold that has been explored for DC programming. Synthetic silica rods have a characteristic wide surface area that can be exploited as a platform for the
controlled release of therapeutics, and their ability to form large pores facilitates cellular
infiltration [103]. For instance, Kim et al. employed an in situ-forming injectable porous
scaffold made from high-aspect-ratio MSRs that substantially enhanced the recruitment
of DCs when compared to low-aspect-ratio MSRs. When loaded with GM-CSF,
CpG-ODN, and model antigen protein OVA, high-aspect-ratio MSRs exhibited
sustained release of the immunomodulators in vitro and prolonged the presentation of
OVA antigen in vivo. Unlike bolus vaccine, this injectable scaffold recruited large numbers of immune cells and induced potent Th1, Th2, and antigen-specific T-cell
responses. The induction of robust humoral and cellular immune responses led to a significant delay in tumor progression in a prophylactic lymphoma model. The MSR scaffold itself may also contribute to adjuvant properties, possibly via NLRP3 inflammasome
activation induced by the internalization of degraded low-molecular-weight silica debris
[104]. Furthermore, Li et al. explored the surface chemistry of MSRs to modulate
immune cell trafficking and activation, by functionalizing MSRs with PEG. PEGmodified MSRs were shown to enhance DC activation by upregulating activation
markers such as CD86 and production of IL-1β in vitro, as well as increased immune cell
trafficking in vivo. This suggests that a simple surface modification could have profound
implications for immune cell infiltration within the scaffold [105]. More recently, Li et al.
designed MSRs by adsorbing PEI on its surface to improve antigen immunogenicity.
The authors demonstrated that the model antigen OVA can be adsorbed on MSRPEI, along with GM-CSF and CpG-ODN. Unlike bare MSR-based scaffold, their
approach (i) enhanced the number of recruited DCs, (ii) improved DC activation, antigen presentation, and trafficking to the lymph node, and ultimately (iii) induced CTL
responses. Their MSR-PEI vaccine was also tested with neoantigens from multiple
tumor models and elicited robust antitumor responses [106]. As previously discussed,
preformed scaffolds can overcome many limitations of in situ-forming gels. Injectable
preformed matrices need to possess shape-memory properties and interconnected
macropores to aid cell trafficking. Generally, hydrogels have small pores in the nanometer
range that do not facilitate cellular infiltration [46,107]. To improve upon this, Verbeke
et al. designed porous alginate-based hydrogel by integrating beads that act as sacrificial
porogens to form large pores. This formulation, when incorporated with GM-CSF,
79Polymeric scaffolds for antitumor immune cell priming
80 Khushbu Bhatt et al.
significantly improved cellular infiltration and, in particular, recruited large numbers of
immature DCs [108]. In lieu of beads, cryogelation is a technique that utilizes ice crystals
as porogens, leading to the formation of an interconnected macroporous network upon
melting. This prevents any toxicity from leftover bead components or the use of toxic
organic solvents. Similar to the PLGA-based implantable scaffold developed by Ali
et al., Bencherif et al. designed an alginate-based cryogel matrix with shape-memory
properties for cancer vaccine applications utilizing GM-CSF as a DC recruiting factor,
CpG-ODN as an adjuvant, and irradiated melanoma tumor cells as an antigen source.
Cell-adhesive RGD peptides were incorporated to promote tumor cell adhesion to
the cryogel scaffold. The resulting cryogel exhibited large, interconnected macropores,
thereby conferring a more favorable environment for cellular infiltration, along with
enhanced mechanical stability, as compared with standard (mesoporous) hydrogels. This
vaccine exhibited sustained release of immunomodulators over 30 days. Furthermore,
robust long-term prophylactic and therapeutic antitumor immune responses were
observed in a murine melanoma tumor model [75,109]. More recently, Shih et al.
improved the injectability and mechanical stability of the alginate-based cryogel scaffolds
by incorporating combined ionic and covalent crosslinking. These tougher cryogel
matrices facilitated injection through a smaller-gauge needle, thereby limiting tissue
damage after injection. When used to formulate cancer vaccines, these improved
cryogel-based scaffolds prevented tumor growth in 80% of mice in a breast cancer model
[71]. Cryogels can also be designed to serve as a cell-responsive platform. For instance,
Koshy et al. described an injectable, cell-responsive, biodegradable gelatin-based cryogel
scaffold encapsulating GM-CSF. Gelatin-based scaffolds inherently possess cell adhesion
properties and can be enzymatically degraded by matrix metalloproteinase (MMP). Upon
injection, these scaffolds elicited controlled release of GM-CSF that led to the infiltration
of immune cells, ultimately leading to MMP-induced cell-mediated degradation of the
cryogel [110].
Overall, injectable polymeric scaffolds can be: (i) easily customized based on the biological need, (ii) fine-tuned for controlled delivery of immunomodulators, and
(iii) combined with other immunotherapies for synergistic tumor rejection.
Additionally, the 3-D porous scaffolds are excellent tools for in situ DC recruitment
and programming (Fig. 2), thus supporting their potential use in therapeutic cancer vaccination. However, the mechanisms by which the polymeric scaffolds encapsulating
GM-CSF and CpG-ODN are involved in DC recruitment, activation, and maturation
have not been thoroughly explored. More studies are needed to dissect the effect of
immunomodulators in conjunction with the polymeric scaffold in exerting antitumor
immunity. Furthermore, it is still unclear whether long-term recruitment and activation
of DCs is beneficial, and the ideal time frame required for optimal DC recruitment and
stimulation needs to be further investigated.
81Polymeric scaffolds for antitumor immune cell priming
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Fig. 2 Schematic illustration of a polymeric scaffold-based cancer vaccine for DC recruitment and
programming. Polymeric matrices are loaded with DC-recruiting and activating molecules. Upon
local administration of the scaffold-based vaccine, there is a sustained release of immunomodulators.
This results in the recruitment of immature DCs within the scaffold where they are presented with
immunostimulatory cues such as tumor cell lysate or antigenic peptides, and adjuvants, leading to
their activation followed by antigen processing and presentation. Subsequently, activated antigenloaded DCs migrate to the draining lymph nodes to prime and trigger T-cell responses. Tumor
antigen-specific CTLs then travel via the blood vessels to reach the tumors, infiltrate them, and ultimately kill cancer cells. A small fraction of T cells is converted to long-lived memory cells, which protect
against tumor recurrence when reencountering the same antigens.
4. Polymeric scaffolds for T-cell activation
While DC-based or DC-targeting cancer vaccines depend on in vivo T-cell
responses, adoptive T-cell therapy depends on ex vivo expansion of T cells that are
subsequently administered into the patients. This approach of circumventing in vivo
T-cell activation offers several benefits such as an environment devoid of immunosuppression and the ability to engineer T cells to produce immunostimulant cytokines as well
82 Khushbu Bhatt et al.
as express antitumor CARs or TCRs [111]. Adoptive cell therapy, however, requires
lymphodepletion and cytokine infusion, as well as generating large numbers (10
11
of functional T cells while preventing activation-induced exhaustion or terminal differentiation, which remains challenging [112]. Thus, strategies to generate high T-cell
numbers in a short time are highly desirable. Currently, commercially available antiCD3- and anti-CD28-coated Dynabeads are widely used to activate and expand
T cells ex vivo. They act as artificial APCs and provide signal 1 of TCR engagement
and signal 2 of costimulation for T-cell activation, and rely on exogenous IL-2 for signal
3. However, these paramagnetic Dynabeads are made with stiff materials and lack ligand
flexibility [111,113]. Even though these cultures provide all three canonical signals
required for T-cell activation, the context in which they provide these cues largely differs
from natural APCs [111,113].
Polymeric scaffolds can provide a fine-tuned 3-D environment to mimic the
physiological activation of T cells in the presence of antigen-presenting cells (Fig. 3).
Rio et al. demonstrated that incorporating a 3-D scaffold system during T-cell activation
with artificial APCs effectively increased T-cell proliferation when compared to standard
2-D culture. The authors attributed the increase in T-cell proliferation to the natural
resemblance of their 3-D structure to the secondary lymphoid organs [114]. More
recently, Rio et al. engineered a heparin-functionalized 3-D PEG-based hydrogel scaffold to more closely resemble lymph node functions. Heparin was used to dock the chemokine CCL21, which promotes T-cell proliferation and migration. The 3-D hydrogel
scaffolds, loaded with CCL21, resulted in increased CD4+ T-cell proliferation and elevated proportions of effector T cells when combined with T cell-stimulating Dynabeads
[115]. Furthermore, Majedi et al. showed that the mechanical stiffness of the microen-
vironment greatly affects T-cell proliferation and function. Three-dimensional porous
alginate-based scaffolds were fabricated to mimic the mechanical properties of the lymph
nodes. They have shown that stiffer scaffolds promoted T-cell activation, expansion,
effector functions, motility, and cell spreading. The authors further demonstrated that
T cells formed larger immune synapses with APCs upon encountering antigens in stiffer
scaffolds, accounting for their enhanced activation and proliferation [116].
Polymeric scaffolds represent an exquisite platform for ex vivo T-cell activation and
expansion since they can be designed to display a high density of T cell-activating signals
and enhanced control over ligand presentation and cytokine release. Furthermore, providing stimulatory cues in a 3-D scaffold more closely emulates the physiological mechanisms involved in T-cell stimulation. For instance, Fadel et al. reported that polymeric
composite scaffolds of carbon nanotubes functionalized with CD28 antibody and
SIINFEKL-MHCI complex increased OVA-specific TCR transgenic (OT-I) CD8+
T-cell proliferation by 1.5-fold, compared to commercially available Dynabeads.
Another 1.5-fold increase in T-cell expansion was mediated by paracrine delivery of
IL-2 from PLGA nanoparticles that were complexed with the carbon nanotubes.
)
83Polymeric scaffolds for antitumor immune cell priming
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Fig. 3 Schematic illustration of polymeric scaffold providing a fine-tuned 3-D microenvironment
for ex vivo T-cell expansion. Polymeric matrices are incorporated with T-cell activation cues such
as CD3 and CD28 antibodies, and cytokines to support T-cell survival and proliferation. Upon seeding
with naïve T cells, the polymeric scaffold provides the required stimulatory signals in a 3-D microenvironment along with the paracrine release of proinflammatory cytokines to mimic physiological T-cell
stimulation. This leads to enhanced T-cell activation and expansion compared with conventional
methods (i.e., Dynabeads). Expanded antigen-specific T cells are simply recovered by mechanical
harvesting and are subsequently injected into the patients to mount an effective antitumor response.
The authors attributed this increased proliferation to a clustered antigen presentation by
carbon nanotube bundles and IL-2 release. This approach achieved T-cell expansion at
therapeutic levels used clinically but with 1000-fold less IL-2 [117]. In a different
approach, Delalat et al. designed a 3-D-printed multilayered lattice to present
costimulatory cues to a large number of cells simultaneously. The lattices were 3-D
printed using a melt electrospinning writing (MEW) technology that facilitated the formation of a highly organized polycaprolactone-based fibrous network and was further
coated with a functionalized polymer to conjugate CD3 and CD28 antibodies. These
functionalized lattices provided enough space for T-cell expansion, facilitated the
exchange of nutrients, and enabled repeated transient interactions to boost T-cell
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