Biomaterials for Recruiting and Activating Endogenous Stem Cells in situ Tissue Regeneration

Abstract

Over the past two decades in situ tissue engineering has emerged as a new approach where biomaterials are used to harness the body’s own stem/progenitor cells to regenerate diseased or injured tissue. Immunomodulatory biomaterials are designed to promote a regenerative environment, recruit resident stem cells to diseased or injured tissue sites, and direct them towards tissue regeneration. This review explores advances gathered from in vitro and in vivo studies on in situ tissue regenerative therapies. Here we also examine the different ways this approach has been incorporated into biomaterial sciences in order to create customized biomaterial products for therapeutic applications in a broad spectrum of tissues and diseases.

Graphical Abstract

in situ Regeneration

The human body is a complex multicellular system of tissues and organs. Upon injury, repair mechanisms are carried out at cellular level to restore the structure and function of damaged tissue [1]. However, in humans the capacity for tissue regeneration and scar-free repair declines with age, partly due to cell-intrinsic signals that diminish the number of functional endogenous stem cells [2, 3]. Chronic and autoimmune diseases, degenerative disorders and cancer also cause harmful effects on the regenerative capacities of stem/progenitor cells by transmitting disruptive signals that hijack the normal physiology of the stem cell niche or the neighboring cell microenvironment [4–6]. Given that the tissue microenvironment is crucial in steering cellular fate towards either homeostasis, repair, or disease, biomaterials have been developed that specifically mimic the structure, composition, and function of the tissue microenvironment to repair or replace damaged or diseased tissues and organs [7, 8].

Tissue engineering methods have traditionally involved the transplantation of these niche-like biomaterials in combination with in vitro-expanded stem cells to a tissue defect site [9, 10]. However, the generation of these cell-laden constructs for clinical use entails multiple, complex steps and poses significant limitations. First, the preparation of cells for transplantation offers numerous challenges, including the difficulty of standardizing tissue culture conditions and reagents, inability to generate a sufficient number of appropriate cells required to populate the biomaterial, and loss of cell viability or phenotype. Once the cell-laden biomaterial is transplanted, another set of obstacles persist, including the potential for immune rejection, tumorigenesis, and pathogen transmission. In face of these risks, the likeliness for these products to pass the Food and Drug Administration (FDA)’s rigorous safety regulatory requirements for clinical application approval are very limited [11].

The emergence of new insights in developmental and stem cell biology relating to the control of stem cell fate, has greatly advanced the field of regenerative medicine by inspiring the creation biomaterials that leverage the innate regenerative potential of resident stem cells to restore the function and structure of damaged or diseased tissues [2]. This cell free approach is known as in situ tissue regeneration, which serves as an alternative to stem cell transplantation and operates by recruiting endogenous stem/progenitor cells to the injured or diseased site and priming them for tissue-specific functions [11]. The absence of transplanted cells in this therapeutic strategy circumvents the critical barriers hindering the clinical use of traditional cell-laden constructs in tissue engineering. Given the lack of in vitro cell expansion and the lack of an exogeneous cell source, in situ tissue regeneration therapy is simple, faces fewer safety and regulatory hurdles, and only relies on the biomaterial’s interaction with local stem/progenitor cells in their natural in vivo microenvironment. Taken together, in situ regeneration represents a more straightforward therapeutic approach to tissue engineering and regenerative medicine. However, unlike cell-laden constructs, the biomaterial designs utilized for in situ regeneration must also encompass properties that recruit stem/progenitor cells as well as properties that support stem/progenitor cell proliferation and differentiation toward mature cell phenotypes.

Adult stem/progenitor cells and their homing factors

Adult stem/progenitor cell populations in tissues and organs has emerged as an attractive source of multipotent cells and a therapeutic target for in situ regenerative therapies, due to their capacity for self-renewal, ability to give rise to mature cell lineages, or, in some cases, travel to distant sites to participate in the repair and regeneration processes of damaged tissues [12, 13]. Stem cell transplantation studies provide essential insights into the cell surface biomarkers that may be potentially targeted by cell-free biomaterials to recruit stem/progenitor cells to injury or disease sites for in situ regeneration (Table 1) [14–29]. Adult stem/progenitor stem cells reside in specialized microenvironment called stem cell niche, that harbors multiple cells such as fibroblasts, endothelial cells, and/or stromal components. The niche and its components tightly regulate the behavior and function of stem cells through direct interactions and/or signaling cues from soluble factors [12].

Table 1.

Adult progenitor/stem cells and putative therapeutic uses

Adult progenitor stem cells	Niche	Surface biomarkers	Cell progeny	Putative clinical application	Refs.
Hematopoietic stem cells (HSCs)	Bone marrow & vessel walls	CD34− or CD34+/CD38−/Thy+/CD90+/CD117−/CD133+/vascular endothelial growth factor receptor 2 (VEGFR2+)	Blood cells; platelets	Autoimmune diseases, anemia, thrombocytopenia, leukemia, aggressive solid tumors	[14–16]
Mesenchymal stem cells (MSCs)		CD49a; CD133; CD45low/med; CD271; CD73; CD105; CD90; CD44	Chondrocytes; cardiac cells; insulin-producing β cells	Cartilage disorders, osteoarthritis; cardiovascular disorders; diabetes	[17, 18, 20]
Endothelial progenitor cells (EPCs)		CD34+; CD133; VEGFR-2/Fetal liver kinase (Flk-1); CD117; Cxc chemokine receptor-4 (CXCR)	Endothelial cells	Vascular disorders	[22]
Neural stem cells (NSCs)	Brain	CD133+/nestin	Brain cells	Nervous system disorder	[24, 25]

During normal physiological conditions, adult stem/progenitor cells maintain their pool through symmetric cell division, but also undergo differentiation via asymmetric division to replenishing local mature cells of the niche. Asymmetric stem cell division gives rise to one daughter stem cell and an intermediate or early transit-amplifying (TA) cell with migratory abilities [12]. At the onset of tissue injury, molecular and physiological changes within the local stem cell niche activate the migration, amplification, and terminal differentiation of TA cells into their respective cell progenies at local or distant sites to repair the damaged tissue [12]. The process of stem cell recruitment by molecular factors from their niche to distant sites, to participate in tissue repair is referred to stem cell homing [30]. Stem cell homing is a process where stem cells are recruited to a site in response to gradients of these chemoattractants by migrating up these gradients and sheltering within injured or pathological tissue areas. The process was initially delineated for transplanted hematopoietic stem cells that migrate from peripheral to bone marrow stem cell niches. In another example, stromal cell-derived factor (SDF-1) is one of the common molecular signals that has been reported to be released after injury to the liver, brain, and heart. High levels of SDF-1 were reported to cause a significant mobilization of bone marrow derived mesenchymal stem cells (BM-MSCs) at injury site for the recovery of the respective tissues [31–33]. To this end, potent homing factors (Table 2) known to recruit and attract stem/progenitor cells duing tissue injury and repair may be incorporated into the biomaterial design to support the recruitment and migration of the desired stem/progenitor cells for in situ regeneration.

Table 2.

Signaling molecules involved in stem cell homing for in situ regeneration

Homing factor	Target cell	Target cell receptor	Clinical application	Refs.
Substance P	BM-MSCs; stromal-like cells	Neurokinin 1 receptor (NK1R); CD29	Bone repair; neovascularization; osteoarthritis; angiogenesis	[36, 37]
Stromal-derived factor 1 (SDF-1)	HSCs; MSCs; EPCs; NPCs	C-X-C chemokine receptor 4 (CXCR4)	Neurogenesis; neural repair; angiogenesis; tendon regeneration; dental pulp repair/regeneration; osteochondral repair	[38–43]
Granulocyte-macrophage colony-stimulating factor (G-CSF)	HSCs; BM-MSCs; EPCs	—	Neurogenesis, neural repair/regeneration; angiogenesis; repair; renal tubular repair; Vascularization	[44–46]
Stem cell factor (SCF)	HSCs	CD34; CD117	Vascular wall repair; dental pulp repair/regeneration; renal tubular repair;	[47, 48]
Monocyte chemotactic proteinl &3 (MCP-1&3)	Angiogenic cells; MSCs	CC-chemokine family	Angiogenesis	[49]
Transforming growth factor p3	BM-MSCs; synovium SCs	—	Cartilage repair	[50]

The role of macrophages in in situ tissue regeneration

The critical role of the immune system in tissue repair following injury is well documented [34, 35]. During the first step of healing, injured tissues signal their need for repair by releasing of a cocktail of chemoattractant signals (Table 2) to reparative cells including stem/progenitor cells and immune cells [36–50]. Although achieving optimum stem cell homing at injury site is an essential step towards tissue repair, the new “home” has to be accommodating to the migrating endogenous stem/progenitor cells to allow their residence. One strategy to ensure successful stem cell residence at the target site is through modulation of immune response to generate reparative signals that promote stem/progenitor cell differentiation. There is mounting evidence pointing to the macrophages’ ability to steer the microenvironment from an inflammatory niche to reparative state via macrophage polarization from M1 macrophages (pro-inflammatory) to M2 macrophages (anti-inflammatory response) [51]. The induction of M2 macrophage phenotypes is induced by the presence of cytokines, such as interleukin-4 (IL-4), IL-10, and IL-13, and supports M2 macrophage reparative functions including phagocytosis, tissue regeneration, and angiogenesis [52]. In contrast, potent inflammatory cytokines, such as tumor necrosis factor alpha (TNF-α), supports M1 macrophages that contribute to the inflammatory and degenerative environment by secreting TNF-α, IL1-α, IL6, ILβ [52]. For example, during degenerative musculoskeletal disease, such as osteoarthritis, sustained breakdown of the extracellular matrix inhibited mesenchymal stem cell homing and differentiation towards cartilage repair [53, 54].

Although much of the current knowledge on macrophage polarization toward the regenerative M2 phenotypes is largely based on the injured or pathological niche, the difficulty of controlling the multicellular and complex in vivo microenvironment underscores the necessity to identify other readily controllable cues supporting M2 macrophages. For example, topographic and architectural cues embedded within the biomaterial may play a role in coaxing macrophage polarization to further support endogenous stem/progenitor cell homing and differentiation for tissue regeneration. Thus an in-depth exploration of the chemical and physical properties of biomaterials that induce macrophage polarization is conducive to robust in situ tissue regeneration.

Designing biomaterials for in situ tissue regeneration

The goal for biomimetic design for in situ regeneration is to create complex matrix systems, based on the composition and structure of the native tissue, that are able to recruit, accommodate, and steer the fate of endogenous stem/progenitor cells towards tissue regeneration without causing harmful inflammation. Following implantation, biomaterials are infiltrated by fibroblasts and macrophages that can either evoke an inflammatory or regenerative response. Biophysical characteristics of biomaterials such as porosity, micro/nanoscale surface patterns, architecture, and stiffness/elasticity have the ability to alter the local microenvironment via cell-biomaterials interactions once implanted, subsequently, influencing the behavior of endogenous stem cells via direct guidance contact or immune-mediated pathway. Designing the chemical and physical properties of the biomaterial is central to controlling immunologic responses, stem/progenitor cell recruitment and stem/progenitor cell fate.

Biomaterial chemical properties critical for in situ regeneration

Biomaterial selection: natural, synthetic, or composite.

The selection of the biomaterial is critical for in situ regeneration, given that biomaterials possess innate biochemical properties can directly impact the host immune response and subsequent stem/progenitor cell recruitment and fate. Biomaterials for in situ tissue regeneration applications can be derived from natural, synthetic, or composite materials. Natural biomaterials are those derived from native tissues, and can represent individual polymeric molecules (Table 3) or as an entire de-cellularized native tissue (Table 4) [55–65]. Synthetic biomaterials can be derived from FDA-approved polymers, such as poly(lactic acid) (PLA), polylactide caprolactone (PCL), polyglycolide (PGA) with excellent biodegradable and biocompatible properties [66]. However, natural biomaterials are relatively more attractive materials for in situ tissue regeneration due to their superior biocompatibility and degradability, endowing them with an immunogenic character that is lacking in many synthetic biomaterials [2]. However, there are several practical approaches known to improve biocompatibility in synthetic materials, such as alteration of their surface hydrophilic/hydrophobic level and surface charge. For example, studies comparing immunomodulatory capabilities between hydrophilic or hydrophobic-coated surfaces, found that hydrophilic-coated surfaces (alkalamine, acrylic acid, 2-methyl-2-oxazoline) promoted albumin adsorption, resulting in the production of anti-inflammatory cytokines. In contrast, hydrophobic surfaces such as 1,7-octadiene-coated surfaces promoted adsorption of immunoglobulins, resulting in the production of pro-inflammatory signals by local macrophages [67]. Furthermore, it was also reported that immune cells responded differently to positively and negatively charged biomaterials. For example, gold nanorods that were coated with negative charged groups such as carboxy (COO⁻) were able to activate inflammasomes, which in turn would trigger pro-inflammatory signaling to a higher degree than positive-charged groups such as amino (NH₃⁺) [68]. These are potential strategies to alter the surface of biomaterials in order to generate chemical cues that evoke proper immune responses.

Table 3.

Natural biomaterials and their applications for in situ regeneration

Biomaterial	Structure/shape	Bioactive molecules	Animal model	Outcome	Refs.
Collagen	Sponge	bFGF/gelatin microspheres	Dog periodontal defect	New vascular and osteogenic tissues & recovery of periodontal ligament	[55]
Collagen	Sponge	—	Rabbit skeletal muscle injury	Myofiber regeneration	[56]
Collagen	Hydrogel	—	Rabbit articular cartilage defect	Recruitment of BM-MSCs for cartilage repair	[57]
Collagen	Hydrogel	—	Rat renal reperfusion/ischemia	Recruitment of renal progenitor/stem cells & recovery of renal function	[58]
Collagen	Porous bilayer	—	Rat diaphragm defect	Promote cellular in growth and alignment & maturation of blood vessels	[59]
Gelatin	Scaffold	IGF-1	Rat skeletal muscle injury	Recruitment & differentiation of Pax7+ progenitor stem cells (Pax7+ cells) for accelerated muscle regeneration	[60]
Gelatin	Hydrogel	rhFGF-2	Mouse bone maxilla defect	rhFGF-2 sustained & controlled release for bone formation	[61]
Hyaluronic acid (HA)	Hydrogel	REG peptide embedded in hydrogel	Mouse cutaneous wound	Accelerated re-epithelialization for skin regeneration	[62]

Table 4.

De-cellularized ECM and their applications for in situ regeneration

Biomaterial	De-cellularization reagents	Animal model	Outcome	Refs.
Porcine esophageal mucosa	1% trypsin/0.05% EDTA/37 °C/1h; 3% Triton X-100/48h	Rat abdmoninal esophagus defect	Oesophagal mucosa reconstruction	[63]
Porcine dermal tissue	2% sodium deoxycholate/10mM HEPES/24h; DNAse overnight/4C	Primate abdominal wall excision repair	Repopulation & maturation of vascular cells; transient immune response	[64]
Porcine urinary bladder matrix	0.1% peracetic acid; 4% ethanol/2h	Dog hemilarynx defect	Re-epithelialization; formation of cartilageous structure; normal phonation; scarred regenerated vocal mucosa	[65]

Although superior in biocompatibility, natural biomaterials are plagued with poor mechanical strength due to rapid degradation once implanted [69]. Mechanical stability of biomaterials provides the necessary framework for cell and tissue infiltration following their in vivo implantation. Therefore, rapid degradation should be avoided, because newly formed tissue requires time before becoming fully-functional. To improve their mechanical integrity, natural biomaterials are often combined with synthetic ones to produce hybrid or composite biomaterials that acquire advantages of both categories (Table 5) [70–77]. Besides the degradation rate of biomaterials, degradation by-products are also important factors to consider. These by-products can be controlled to further modulate the microenvironment for stem cell homing and/or evoke proper immune responses. Many readily-biodegradable synthetic biomaterials such as poly-lactic-glycolic acid (PLGA) and poly-l-lactide PLLA, although non-cytoxic, their acidic degradation by-products have been widely recognized as detrimental to cell mobilization and angiogenesis as indicated in a study by Sung and colleagues in which they reported a lower density of migratory cells and poor vascularization at implantation site where PLGA scaffolds were implanted compared to PCL scaffolds. Not only PLGA scaffolds were found to degrade faster, but also the degradation by-products increased the acidity in the environment, thus, causing less to migrate and proliferate [78]. On the other hand, degradation by-products like bioactive ions can improve tissue regeneration. For example, the degradation of microporous hydroxyapatite scaffolds released phosphate, calcium, and magnesium ions in rabbit bone defects and promoted the formation of new blood vessels [79]. Furthermore, in a different study, release of magnesium and calcium ions from silicate bioceramic scaffolds caused positive immunomodulatory effects conducive to tissue regeneration, by inducing macrophage anti-inflammatory response [80]. Natural biomaterials such as collagen and chitosan, on the other hand, are naturally susceptible to proteolytic or enzymatic degradation and remodeling by cells, and their by-products are less likely to cause cytotoxic effects. Furthermore, some of these natural by-products directly contribute to the regeneration of tissues whether by inducing cell mobilization or modulating gene expression [81]Bioactive homing and differentiation cues. Cytokines, growth factors, and synthetic agents such as Wnt inhibitor sclerostin, resolvin D1, and sphingosine phosphate agonist SEW2871 (Table 6) promote stem cell recruitment and tissue regeneration [36, 82, 83]. However, in many cases, administration alone of these biomolecules does not necessarily lead to appropriate cellular responses. For example, delivery of SDF-1, a potent stem cell homing factor, under normal conditions is susceptible for failure as SDF-1 is naturally degraded by endogenous CD26/dipeptidyl peptidase IV (DDP-IV) in vivo. To prevent premature SDF-1 degradation in a non-invasive manner, DDP-IV inhibitor such as parathyroid hormone is administered first, followed by SDF-1 administration [84].

Table 5.

Composite biomaterials and their applications for in situ regeneration

Biomaterial	Biomaterial structure	Synthesis method	Bioactive molecule	Animal model	Outcome	Refs.
PCL/HA/chitosan	HA/chitosan-coated PCL fibrous scaffold	Electrospinning; layer by layer stacking	—	Rabbit ligament injury	Fibroblast infiltration & proliferation; poor mechanical strength	[70]
PCL/β-tricalcium phosphate/collagen	Collagen I-coated PCL/β-tricalcium phosphate scaffold	Selective laser sintering	—	Mice bone defect	Enhanced osteogenic & differentiation of ADSCs; bone & vascular tissue formation	[71]
PCL/HA	HA-functionalized PCL nanofibrous scaffold	Electro-spinning	Substance P	Mice calvarial bone defect	Recruitment of bone marrow-derived stem cells; bone formation	[72]
Methacrylated/HA/PLGA	PLGA-reinforced MA/HA macroporous scaffold	Directional cooling; Freeze-drying	—	Rabbit knee cartilage defect	Regeneration of cartilage & bone; anti-inflammatory response	[73]
HA/methyl cellulose	Hydro-gel	Freeze-drying	Erythropoietin	Mouse brain injury	Attenuated inflammatory response; migration & differentiation of NSCs	[74]
PGA/HA	Felt implant	Freeze-drying	Allogenic serum	Rabbit IVD injury	Disc regeneration; mechanically stable	[75]
Poly-4-hydroxy-butyrate/gelatin	scaffold	Jet spinning	—	Heart valve injury	Scaffold infiltration by valvular interstitial cells; valvular tissue growth	[76]
PCL/fibrin	Fibrin-infused PCL scaffold	Electro-spinning	MCP-1	Rat vascular defect	Smooth muscle formation; fully regenerated endothelium	[77]

Table 6.

Biomolecules with immunomodulatory and/or stem cell homing and differentiation properties for in situ regeneration

Biomolecules	Clinical application	Refs.
Substance P + heparin	Vascular regeneration	[36]
SDF-1 + sphingosine phosphate agonist (SEW2871)	Skin wound closure	[81]
Lipoxin A4 (LxA4) + resolvin D1 (RvD1)	M2 macrophage polarization	[83]
Interferon-gamma (IFNy) + interleukin 4 (IL-4)	Bone formation	[84]
Bone morphogenic protein(BMP): BMP-2 + substance P; BMP-2 + SDF-1; BMP-7 + SDF-1	Bone formation; periodontal regeneration	[85–87]

Introduction of these biomolecules to biomaterials not only prevent premature degradation and delivery to non-target tissues, but also increases their visibility to endogenous cells [36, 38, 85–89]. For example, encapsulation of SDF-1 in a chitosan-based nanoparticle allowed a sustained release of SDF-1 for up to 7 days, resulting in significant recruitment of MSCs via activation of P13K/Akt pathway [90]. Although delivery of a single homing factor has been successful in the recruitment of endogenous stem cells and promote their bioactivity, tissue regeneration will likely benefit from presentation and release of multiple therapeutic agents in a spatiotemporal pattern. For example, SDF-1 and BMP-2 were co-delivered by nano-hydroxyapatite scaffold to recruit BM-MSCs and stimulate osteogenic differentiation, respectively, for bone formation in critical-size rat bone defects (Fig. 1) [88]. To ensure rapid BM-MSCs homing, SDF-1 was adsorbed on the surface of the scaffold and released in a burst release manner. To ensure sustained release of BMP-2 for BM-MSCs osteogenic differentiation and long-term bone formation, BMP-2 was loaded inside silk fibroin microspheres and introduced in the inner structure of the hydroxyapatite scaffold. BMP-2 was released was sustained for up to 3 weeks [88]. In other cases, immunomodulatory molecule and a homing factor can be co-delivered to ensure enhanced endogenous stem cell recruitment via modulation macrophage M2 polarization. For example, heparin and substance P were covalently immobilized on the surface of polylactide caprolactone scaffold to suppress thrombogenic responses via macrophage M2 polarization, and recruit BM-MSCs, respectively, for vascular regeneration in subcutaneous rat defects [91].

Figure 1: Controlled delivery of multiple biomolecules for multifunctional purposes.

Physical adsorption of SDF-1 within hydroxyapatite scaffold allows burst release, while BMP-2 encapsulation inside silk fibroin microsphere provides a sustained release critical for bone repair.

Upon stem/progenitor cells migrating to the biomaterials, biochemical cues can also be incorporated into biomaterial design to promote cell adhesion (Fig. 2A–D). Adhesion proteins such as collagen, fibronectin, laminin, and vitronectin are popular choices of biomaterial-coating to promote stem cell adhesion while reducing potential inflammatory responses caused by synthetic biomaterials (Fig. 2A) [92]. These cell-binding proteins are present in native extracellular matrix and determine cell shape, function, and tissue integrity by binding to cell receptors such as integrin, thus, facilitating transmission of biochemical and biomechanical cues to the cell. Besides adhesion proteins, biomaterials can also be conjugated with protein fragments or peptide motifs (Fig. 2B). For example, incorporation of laminin-derived peptide tyrosine-isoleucin-glycine-serine-arginine (YIGSR) into tunable polyethylene glycol (PEG) hydrogels, induced formation of massive cell aggregates [93]. Whereas, incorporation of collagen-derived integrin-binding site GFOGER in PEG gels enhanced adhesion and osteogenic definition of BM-MSCs than arginine-glycine-aspartate (RGD) peptide, a common adhesion binding site with no biological specificity [94]. Similarly, transplantation of GFOGER-functionalized hydrogels into critical-size murine bone defect significantly increased vascular and bone volume than RGD-functionalized hydrogels even in the absence of vascular endothelial growth factor [95].

Figure 2: General approaches to functionalization of biomaterials with signaling molecules for delivery of biological cues.

A. Surface coating of biomaterials with adhesive molecules such as fibronectin, laminin, and collagen. B. Biomaterial surface presentation of adhesion peptides such as arginine-glycine-aspartate (RGD) and proline-histidine-serine-arginine-asparagine (PHSRN) peptides via a double-branched platform. C. Sequestration of biomolecule within biomaterial structure via physical interactions. D. Microsphere encapsulation of biomolecules within biomaterial structure.

Simultaneous presentation of more than cell-binding peptides increases the magnitude of adhesion-mediated cell signaling. For example, co-presentation of RGD and its synergy motif proline-histidine-serine-arginine-asparagine (PHSRN) peptides enhances signaling responses that stimulate hMSCs osteogenic differentiation in vitro and bone formation in rat calvarial defects [96]. Furthermore, spatial arrangement and functionalization strategies of these biomolecules have profound influence on stem cell adhesion. Simultaneous presentation of RGD and PHSRN peptides was achieved via covalent anchoring on titanium surface using a double-branched platform (Fig. 2B).

Physical properties of biomaterials critical for in situ regeneration

Porosity.

Generally, microporous structures offer many advantages over their non-porous counterparts, such as: enhanced cell infiltration and tissue ingrowth, vascularization, sequestration and presentation of signaling molecules, nutrients diffusion and waste removal [97]. In addition, pore size is an important parameter in load-bearing tissues such as the bone, to ensure proper balance between endogenous cell infiltration and engraftment, and maintenance of structural integrity of the biomaterial. In general, small pore structures (200–300 μm) present a better environment for cell seeding but with reduced cell proliferation and differentiation, larger pore structures are more inclined to mechanical failure [98]. Moreover, porous structures support tissue ingrowth and vascularization by providing an interconnected pore networks that enhance cell migration and bioactivity [99]. Interestingly, porous structures may also present a more pro-regenerative environment than non-porous as reported in a study where a 34 μm microporous polycaprolactone scaffold was able to recruit a significant number of macrophages to the surface of the scaffold, and induce significant M2 macrophage polarization [97].

In summary, the interconnected pore network within porous structures promotes endogenous cell infiltration and migration, thus facilitating tissue ingrowth and vascularization. In the absence of homing factors, porous structures may provide stem cell-recruiting capabilities by stimulating M2 macrophage polarization.

Micro/nanopatterns.

Topological features such as micro/nanopatterns may have an effect on endogenous cell-recruitment capabilities. Once in contact with cells, micropatterned structures may impact cell fate and promote or suppress cell adhesion, which in turn will determine the stem cell morphology and by extension determine stem cell progeny. For example, a system of micro-grooved structures on polyimide surfaces using a photolithography technique (Fig.3A) was used to study the impact of the surface topography on MSCs growth and differentiation. MSCs grew parallelly to the main axis of the grooves. On the micro-grooved patterns with thinner ridges (~ 2 μm thick), MSCs acquired an elongated shape and expressed osteogenic markers. However, on thicker ridges (~ 15 μm thick), MSCs were round and expressed adipogenic markers. Interestingly, when MSCs were seeded on nanogrooved surfaces of polyimide chip (600 μm diameter, 650 nm periodicity, and 200 nm deep grooves) (Fig. 3B), MSCs adopted an elongated shape with no lineage-specific tendency, as they were able to express adipogenic (adiponectin, peroxisome proliferator-activated receptor gamma, and fatty acid binding protein 4) and osteogenic genes (secreted protein acidic and rich in cysteine, secreted phosphoprotein 1, and bone gamma-carboxyglutamic acid-containing protein) when treated in the respective differentiation media [100].

Figure 3: Surface patterning an its influence on stem cell differentiation.

A. Schematic cross of grooves and ridges on polyimide substrate. B. Schematic representation of nanogroove structures with peak-to-peak distance (periodicity) of 650 nm on polyimide surface. C. Alterations in biomaterial surface patterning modulates mesenchymal stem cell differentiation toward either adipocytes or osteoblasts. Long fibronectin pattern islands promote elongation of MSCs and osteoblast differentiation, short fibronectin islands promote MSCs round shape and adipocyte differentiation.

Furthermore, surface patterns of certain size were found to contribute to macrophage polarization towards pro-regeneration. M2 polarization (+ M2 markers CD206, IL-10, and arginase 1) were significantly upregulated when macrophages were cultured on a large-size (~ 12 – 36 μm diameter) rather than smaller size (>12 μm diameter) micropatterned surface. In the same study, macrophage M2 polarization was reported to cause angiogenic and osteogenic effects on human umbilical vascular endothelial cells (HUVECs) and BM-MSCs, respectively, via expression of angiogenic (vascular endothelial growth factors, basic fibroblast growth factor, endothelial nitric-oxide synthase, and osteogenic (bone morphogenic protein-2 (BMP-2), collagen-1 (COL-1), and runt-related transcription factor 2 (RUNX-2)), to promote better bone regeneration [101].

Nanoscale topographies have the potential to affect stem cell fate due to the analogous size to cell receptors through contact guidance [101]. However, these patterns do not cause same effects on the cells. The size and the distance between these structures (pattern islands) are the most determining factors on the cell fate. Generally, the smaller the pattern island, the rounder the cell becomes, and the larger the island, the flatter the cell becomes (Fig. 3C) [102].

Fiber orientation.

Electospinning technique allows fabrication of nanofibrous meshes of specific diameter and orientation to recapitulate the structural and functional characteristics of native connective and load-bearing tissues such as ligaments, cartilage, and bone. For instance, the impact of aligned versus randomly-orientated poly(3-hydroxybutyrate-co-3-hydroxyvalerate) nanofibers was determined in bone formation using a rabbit bone defect model. Although no observable difference in bone growth was found, there was a significant difference in mechanical strength, where aligned fibers promoted bone growth of similar elastic modulus to healthy rabbits than did the randomly-oriented fibers [103]. Furthermore, mechanistic investigations on the effect of fiber orientation on MSCs differentiation, revealed that aligned poly-L-lactide acid (PLLA) nanofibers caused osteogenic commitment of MSCs, through cross-regulation between microRNAs and log non-coding RNAs [104]. Assessment of MSCs differentiation in response to PLLA nanofibers orientation was also conducted in rat tendon injury model. It was found that randomly-oriented fibers triggered MSCs osteogenic differentiation by alkaline phosphatase expression, whereas aligned fibers enhanced MSCs tenogenic differentiation by scleraxis gene expression [105].

Fiber size.

The fiber size is a critical biomaterial physical parameter that can direct stem cell differentiation. The impact of PLLA fiber diameter on annulus fibrosus stem cells (AFSCs) differentiation was assessed for potential application in vertebral disc regeneration. Results showed that the AFSCs, which are normally of fibrocartilaginous origin and were extracted from New Zealand white rabbits, acquired a round morphology and expressed abundant collagen-2 and aggrecan markers on small-size fibers (0.55 ± 0.09 μm), whereas on large-size fibers (3.29 ± 0.53 μm), AFSCs were more spindle-like (fibroblastic) and expressed abundant COL-1 markers [106]. These results are in agreement with what is known about the development of native annulus fibrosus (AF) cells within intervertebral tissue compartments. Phenotype of AF cells gradually changes from fibroblast-like in the dense collagen fibrils-packed outer zone to chondrocyte-like in the less-packed collagen bundles inner zone [107].

In summary, nanofibrous matrices have the potential to direct stem cell fate towards specific cell lineage by guidance contact, depending on the size and orientation of the nanofibers. However, mechanical strength of these nanofibers has to match to the native tissues in order to generate functional tissues.

Stiffness and elasticity.

Stem cells can sense mechanical cues from the microenvironment and transduce them into internal forces to restructure and rearrange of the cytoskeleton network, leading to stem cell differentiation by mechanotransduction [102, 108]. It is important to achieve proper rigidity or elasticity within a fibrous biomaterial, to ensure regeneration of functional tissues. For example, naïve MSCs were cultured under the same 2D culture conditions on collagen-coated gels of different rigidities [108]. Results showed MSCs acquired distinct morphologies leading to different progenies. Elastic matrices (0.1–1 kPa) promoted a filapodial morphology leading to neurogenic differentiation, medium stiffness matrices (8–17 kPa) promoted fibroblast-like shape leading to myogenic differentiation, and stiffer matrices (25–40 kPa) promoted round shape leading to osteogenic differentiation. However, these differences in MSCs fate were not reproduced within 3D culture conditions, where MSCs were encapsulated inside alginate/agarose hydrogels of different stiffnesses [109]. Instead, matrix stiffness was demonstrated to control the molecular interface between cells and matrix via integrin binding, thus controlling cell bioactivity.

Mechanical properties can also have a direct effect on immune response, which in turn may determine stem cell fate. Generally, stiffer matrices (~840 kPa) tend to trigger pro-inflammatory responses compared to softer matrices (~130 kPa) that tend to suppress those immune responses [110]. Given the ability of immune cells especially macrophages to sense a wide range of biomaterial stiffness and modulate their immune responses accordingly, immune response can also be manipulated via presentation of immunomodulatory molecules using biomaterials, to reverse an undesirable immune response. For example, low-stiffness gelatin-based maintained self-renewal of BM-MSCs, and a high-stiffness hydrogel promoted osteogenic differentiation [111]. Unfortunately, the high-stiffness gel also polarized macrophages towards the pro-inflammatory response, consequently, hindering BM-MSC osteogenic differentiation. To restore BM-MSC osteogenic potential, the high-stiffness hydrogel was loaded with IL-4 to induce macrophage M2 polarization, and with SFD-1 to recruit endogenous cell homing for rat periodontal repair (Fig. 4) [112]. Results showed improved tissue recovery in periodontal defects that were treated with high-stiffness hydrogel containing IL-4 and SDF-1. Although this is approach generated promising outcomes, further research on immunomodulatory capabilities of biomaterials is critical.

Figure 4: Biomaterials can provide multimodal mechanical and biological cues to synergistically induce regeneration.

Hydrogels with high stiffness that are embedded with SDF-1α and IL-4 promote both an anti-inflammatory response and bone regeneration in vivo.

Although the underlying mechanisms by which topographic and mechanical cues exert influence on stem cell behavior are not well understood, numerous studies have shown that these mechanisms may involve rearrangement of the cytoskeleton network via clustering of cell receptors, such as integrins once they bind to the surface of the biomaterial, which in turn may activate Wnt, Yes-associated protein, and c-Jun N-terminal kinase signaling pathways and cause changes in gene expression [102, 113, 114]. Despite the undeniable effects of biophysical cues (topographic and mechanical cues) on stem cell behaviors for tissue regeneration purposes, these factors alone are not enough. Instead, synergistic effects from combining biophysical and biochemical factors will better mimic in vivo microenvironment to improve tissue regeneration outcome.

Tissue regeneration by cell reprogramming

In diseased tissues with limited regenerative capabilities due to either an abundance of terminally differentiated cells or a shortage of proliferating cells, cell reprogramming is a very attractive tissue regenerative approach. Cell reprogramming includes deriving the desired cells from either direct cell reprogramming or from induced pluripotent stem cells (iPSCs) [115]. Given the complexity and difficulty of using iPSCs, such as the complexity, length, and cost of cell processing and potential tumorgenicity [116], we focus on direct cell reprogramming and cell reprogramming of injured/damaged cells into repaired cells. There are five plausible cell programming approaches: 1) overexpression of lineage-determining transcription factor; 2) gene silencing by RNA interference; 3), stimulation of protein translation through mRNA delivery; 4) gene editing using clustered regularly interspaced short palindromic repeats (CRISPR) and CRISPR-associated protein 9 (Cas9); 5) epigenetic modifications via biophysical and biochemical cues. While cell reprogramming using viral vectors is the most prominent strategy because of its high delivery and integration efficiency, there are serious safety concerns associated with this approach, such as tumor development and size limitation [117]. As a promising alternative, biomaterials have been utilized in different cell reprogramming approaches to induce genetic or epigenetic modifications.

Overexpression of lineage-determining transcription factors.

Transcription factors regulate gene expression by controlling the rate of gene transcription and/or inducing upregulation/downregulation of genes [118]. By regulating transcription factors, it is possible to give cells new characteristics. Therefore, exogenous intracellular delivery of lineage-specific transcription factors is one way to induce differentiation and reprogram cell identity. Common biomaterial-based delivery systems are injectable hydrogels, electrospun fiber scaffolds, and nanoparticles. For example, a polyethylene glycol-based nanocapsule was used to encapsulate and deliver myoblast determining protein 1 (MyoD1) inside the nuclei of myoblast cells [119]. Delivered MyoD1 was able to drive myoblast differentiation and phenotype change into a multinucleated myotubes capable to give rise to skeletal, cardiac, and smooth-muscle cell. Although nanoparticles are promising candidates for intracellular delivery in terms of high stability, biocompatibility and loading efficiency, they generally have lower yield than active targeting [118]. Furthermore, selection of transcription factors for cellular reprogramming is challenging as it is based on a trial-and-error approach [120, 121]. Gene silencing by RNA interference. Intracellular delivery of microRNAs (miRNA) or small interfering RNAs (siRNA) can cause silencing of specific genes by repression of translation or degradation of mRNA once they bind to the host mRNA [122, 123]. Although cell reprogramming by RNA interference is widely adopted in treatment for cancer and genetic diseases, this approach can also be applicable in regenerative medicine. For instance, miR-222 and neurotrophin-3 growth factor (NT-3) were both incorporated inside a composite collagen-based hydrogel and were delivered in rat cervical incision to induce axon regeneration [124]. The addition of miR-222 to the NT-3-based therapy caused significant increase in neuronal growth, because miR-222 directly inhibits multiple neurogenesis-inhibitory genes, such as phosphatase and tensin homolog. Thus, the potency of miRNAs to target and control multiple genes concurrently makes them ideal therapeutic molecules for injuries involving a plethora of dysregulated genes [125]. While siRNAs are more popular for use in cancer treatment, siRNAs are not as widely used in regenerative medicine, most likely due to their highly specificity to a single mRNA target and limited regulatory spectrum [126].

Stimulation of protein translation by mRNA delivery.

Unlike other cell programming approaches that require nuclear entry, mRNA-based therapy is cytosolic and potentially less harmful considering the low likelihood of genome integration. Once inside the cytosol, mRNA reaches the translational machinery for expression of its protein products [127]. As a result, this approach is more likely to generate high transfection efficiency and is applicable to non-dividing or slow-diving cells [128]. Furthermore, encapsulation of mRNA in polymeric nanoparticles significantly improves its stability and reduces its immunogenicity. For example a recent study demonstrated improved ocular repair by enhanced protein expression in mice that were injected with mRNA-loaded lipid nanoparticles than in those treated with plasmid DNA [129]. While still in its infancy, mRNA-based therapy for tissue regeneration presents significant potential.

Gene editing using CRISPR-Cas9.

Cell fate and identity can also be modulated via gene editing using engineered nucleases such as CRISPR that act as molecular scissors, introducing site-specific breaks at selected genome locations [130]. These breaks can either be repaired via non-homologous end joining (gene silencing), or via homology-directed repair (gene activation). CRISPR-Cas9 is a promising approach for in situ regeneration because of its uncomplex design, high efficiency, widespread application to different cell types, and minimum off-target outcomes. Nanoparticles are excellent delivery system in this gene editing approach. For example, gold nanoparticles decorated with densely packed DNA to load guide RNA, donor DNA, and Cas9 were injected in mouse model of Duchenne muscular dystrophy [131]. Corrected dystrophin gene mutation was validated through expression of dystrophin protein. However, editing efficacy was very low and resulted in partial rescue of muscle function. In spite of this outcome, there is strong potential application of nanomaterials and biomaterials in general, in gene editing for tissue regeneration.

Epigenetic modifications by biophysical and biochemical cues.

Epigenetic modifications directly alter chromatin structure and make DNA more accessible for transcription to regulate gene expression and cell identity [132]. Popular epigenetic modifications involve DNA methylation and histone post-translation modifications. The advantage of this approach lies in its reversible nature. A biomaterial’s biophysical properties such as mechanical stiffness, surface patterning/topography have also shown to cause epigenetic changes in cells. For example, micro-grooved or nanofibrous scaffolds showed higher cell reprogramming efficacy than smoother surfaces. This is attributed to the opening of chromatin structure of seeded cells caused by rough surfaces, thus facilitating interactions between DNA and transcription machinery [133]. Furthermore, these nanopatterned surfaces also showed to have influence on DNA methylation pattern of embryonic stem cells (ESCs), thus directing ESCs towards either stemness or differentiation [134]. Synergistic effects of combining biophysical cues provided by the structure of the biomaterial, and biochemical cues released from epigenetic regulating molecules such as gemcitabine and decitabine, are more effective in stimulating and enhancing tissue regeneration outcomes [135].

Commercially available biomaterials for in situ regeneration

Recent progress in regenerative medicine has led to the development of novel therapeutic approaches that have been translated into clinical applications, such as INFUSE Bone Graft for orthopedic or dental applications, NeuraGen/Neurotube for nerve repair, and GORE-TEX vascular grafts, while many others are a few steps away from bench to clinical translation. These novel therapeutic products were produced using some of the approaches that were discussed in this review of tuning biomaterials properties and combined them with bioactive molecules to enhance endogenous cell responses. These products provide cell-free systems that eliminate the need for exogenous stem cell transplantation that are otherwise laced with many challenges. Progress in biotechnology has facilitated development of recombinant proteins with active signals that can be incorporated with biomaterial structures to enhance biomaterial performance in the restoration of a regeneration-conducive environment. Furthermore, micro- and nanoscale biofabrication strategies has advanced the design of biomaterials structures at resolutions that were previously unattainable, with tunable properties for multifunctional applications.

INFUSE Bone graft is produced from a combination of recombinant human BMP-2 (rhBMP-2) on absorbable collagen 1 sponge [136]. In this system, collagen sponge serves as a delivery system for rhBMP-2 to ensure a controlled and sustained release, thus mitigating potent osteogenic effect. Although, rhBMP-2 was found to be highly osteo-inductive in preclinical studies as early as the 1990s, INFUSE Bone graft (Medtronic Spinal and Biologics) was approved for clinical applications a decade later. Nevertheless, this bone graft was approved for multiple clinical applications: In 2002, for interbody spinal fusion; in 2004, for open tibial fracture; and in 2007, for oral maxillofacial application. Clinical studies in these three categories showed that INFUSE Bone grafts produced if not significantly superior outcomes than autologous grafts (spinal fusion), they performed as well as autologous grafts [136]. Considering severe complications associated with harvesting of autologous bone grafts, these clinical studies demonstrate that INFUSE Bone grafts are among the best bone graft substitutes.

Among several FDA-approved nerve guidance conduits (NGC) are NeuraGen, a collagen 1-based NGC, and Neurotube, a polyglycolic acid-based NGC, for critical-size nerve defects. Given limited surgical procedures for large-size nerve defects, critical-size nerve repair relies on clinical applications of graft materials [137]. NeuraGen, an Integra Life Science Corporation product, was the first semi-permeable type I collagen NGC to be approved by the FDA in 2001. As a major component of the extracellular matrix, collagen offers superior biocompatibility and enhanced cell migration compared to synthetic materials. The fibrillar structure of collagen is maintained throughout the manufacturing process of NeuraGen, permitting a biocompatible tubular matrix with sufficient mechanical strength, defined permeability, and a controlled rate of resorption [138]. NeuraGen is not expected to be fully resorbed within 4 years post implantation [139]. Various clinical studies using NeuraGen for nerve repair in medium to large-size nerve defects (5–18 mm) were conducted following its approval by the FDA. Overall, NeuraGen produced satisfactory results of nerve sensory recovery and low-level post-operative pain [140–142]. Like NeuraGen, Neurotube was the first FDA-approved NGC of the synthetic bioresorbable family in 1999. Unlike NeuraGen, Neurotube has excellent degradability due to flexible mechanical properties, can be applied in larger-size defects up to 3 cm, and has the most clinical available data for review of its safety and efficacy among all FDA-approved devices. Among Neurotube major disadvantages are high degradation rate, acidic degradation byproducts, and low solubility [137].

GORE-TEX vascular grafts are among the many polytetrafluoroethylene-(PTFE) based grafts that have been widely used for hemodialysis access for many years. Unlike other PTFE-based grafts, GORE-TEX grafts (Gore & Associates Inc.) are 100% covered with a thin, porous, and expendable PTFE film, leading to improved handling characteristics [143]. Although GORE-TEX vascular grafts show higher patency rates than other PTFE-based vascular grafts, those patency differences do not necessarily translate into significant differences in vascular repair [143].

Summary and Future Directions

In this review, we discussed adult stem/progenitor cells, localized in niches of adult tissues and organs, that can be mobilized to local or distant injury sites to participate in tissue repair and regeneration. Stem cell recruitment also referred to as stem cell homing, represents an innovative strategy in regenerative medicine to enhance regenerative capabilities of therapeutic solutions. The potential therapeutic solutions discussed in this review for in situ application, involve incorporation of well-selected signaling molecules into natural and synthetic biomaterials for coaxing endogenous stem cell homing by contact guidance or via immune-mediated pathway by macrophage polarization. This cell-free transplantation approach circumvents regulatory-related issues. However, this approach may not be suitable for tissues whereby stem cell number has declined or stem cell dysfunction.

For a long time, mechanisms through which allogenic decellularized tissue grafts have been able to recruit endogenous stem cells and orchestrating regenerative processes in similar manner as autologous grafts in the absence of cells, has been elusive. However, recent findings have revealed some of the unique properties of the ECM that play crucial roles in tissue regeneration such as providing homing signals to recruit local and distant endogenous cells [144]. As a result, tissue decellularization has become an integral part of regenerative medicine, as one of tissue engineering methods that is utilized to produce ECM-derived biomaterials that preserve the complex 3D structure, biochemical and biomechanical properties of native tissues that are crucial for stem cell homing and accommodation.

Significant progress in micro/nanoscale biofabrication has been applied in engineering biomaterials with complex structures, multiscale architectures, and topographic features to exploit the body’s ability to heal and regenerate. Although results from these pre-clinical studies are promising, further research on scaling-up these potential therapeutic solutions to recapitulate as many characteristics of native tissues, developing appropriate animal models that mimic clinical conditions, and approval of standard assessment methods are needed to ensure efficacy, reliability, and safety of these therapies before they can ultimately be used in clinical studies.

Statement of Significance.

Biomaterials can be designed to recruit stem cells and coordinate their behavior and function towards the restoration or replacement of damaged or diseased tissues in a process known as in situ tissue regeneration. Advanced biomaterial constructs with precise structure, composition, mechanical, and physical properties can be transplanted to tissue site and exploit local stem cells and their micro-environment to promote tissue regeneration. In the absence of cells, we explore the critical immunomodulatory, chemical and physical properties to consider in material design and choice. The application of biomaterials for in situ tissue regeneration has the potential to address a broad range of injuries and diseases.