Tissue Engineering and Applications for Treatment of Myocardial Infarction
Tissue engineering is an interdisciplinary field that applies the principles of engineering and life sciences for the creation of biologically active scaffolds that regenerate, restore, maintain or improve tissue function or a whole organ. Many new developments have been made in the field of tissue engineering over the last two decades and, in recent years we have seen a rising number of tissue engineered products in commercial use a variety of different cell types have been used such as stem cells and progenitor cells. An assortment of different natural and synthetic biomaterials have also been utilized as vehicles for cellular delivery or as acellular constructs. This chapter discusses the history, principles and design strategy for tissue engineering therapies with a specific focus on cardiac tissue engineering approaches.
B. Introduction to Tissue Engineering
This book chapter will provide an overview of tissue engineering with an emphasis on cardiac tissue engineering approaches for the treatment of myocardial infarction (MI) and heart failure (HF). In particular, the chapter will outline injectable and patch-based biomaterial scaffolds for cellular delivery as well as for acellular treatments.
History of tissue engineering
Tissue engineering is a relatively young, interdisciplinary field of science intersecting medicine, biology and engineering. The term ‘tissue engineering’ was first coined by Y.C. Fung from the University of California, San Diego in 1985 (1). However, the field was formally defined at the first scientific meeting discussing tissue engineering approaches in 1998 as “the application of the principles and methods of engineering and the life sciences toward the fundamental understanding of structure/function relationships in normal and pathological mammalian tissues and the development of biological substitutes to restore, maintain, or improve functions” (2).
Tissue engineering gained further recognition through the seminal article published by Langer and Vacanti in Science in 1993 (3). They described tissue engineering as a process in which biological and engineering principles are used to bring together cells, biological molecules and scaffolds to augment or replace damaged tissue (3). A marked moment in the field of tissue engineering, was the landmark study by Langer et al., in which a polymer template in the shape of an ear was seeded with chondrocytes and subcutaneously implanted on the back of the mouse (4). This article received widespread publicity putting tissue engineering on the map, while at the same time inspiring and educating researchers and the common public alike of potential life-changing advancements the field of tissue engineering could bring.
Traditionally, tissue engineering utilized cell-seeded constructs that were cultured in a bioreactor and subsequently implanted into the body. It was thought that the cells would proliferate in the scaffold and replace it by secreting extracellular matrix (ECM) components as the original scaffold material degraded. However, more recently many approaches are targeted towards in situ engineering of tissue with or without cells. In the last 20 years, many great advancements have been made in the field of tissue engineering, with several products on the market and in clinical trials (5,6).
Principles of tissue engineering
The “tissue engineering triad” which consists of cells, scaffolds and growth factors/biological molecules, is commonly used to conceptualize tissue engineering (Fig. 1). Tissue engineering approaches involve at least two of the three constituents. However, biomaterial scaffolds alone are often considered to fall under the tissue engineering umbrella as these biomaterials can interact and be infiltrated by host cells, thus adding the cellular component. Each individual constituent of the triad will be discussed in more detail in the subsequent sections. For a more in-depth review many excellent resources are available (3,7,8).
A variety of different cell types have been utilized for tissue engineering applications depending on the target organ (Fig. 2). They may be from an autologous, allogeneic or, in rare cases, a xenogeneic source. The cells can also have the same origin as the intended organ, for example cardiomyocytes have been injected for treatment of myocardial infarction. However, more recently, embryonic stem cells and progenitor cells have been investigated, since with the correct cues they have the potential to differentiate into an organ specific lineage.
Timeline of the progress and different cell types utilized for cell transplantation in cardiac regenerative therapies. Reproduced from (98) with permission from Elsevier.
Biomaterials play a key role in tissue engineering, both as a delivery vehicle for cells and growth factors or as acellular matrices. These materials can provide the much needed biocompatible, biodegradable and biomimetic environment necessary for tissue engineering applications. These scaffolds can allow for tissue growth, promote cell adhesion, migration and proliferation as well as provide mechanical strength. The original inspiration for these materials was as a mimic of the natural ECM components in the body such as collagen, elastin, laminin and hyaluronic acid. Other biologically derived materials such as fibrin, alginate, chitosan and gelatin have also been widely and successfully utilized. As the field evolves, many synthetic scaffolds are also being utilized for this purpose. Though they do not have the benefits of inherent bioactivity, they can circumvent some of the issues seen with biologically derived biomaterials by providing tunability of degradation time, stiffness, and porosity to better target the desired application.
Growth factors/ biological molecules
Growth factors can be delivered with or without a scaffold to promote survival of administered cells or to elicit a biological response from the tissue. Vascularization of tissue-engineered scaffolds can be a challenge and is an essential component to the survival and integration of engineered tissue. Therefore, angiogenesis inducing growth factors such as vascular endothelial growth factor (VEGF) and basic fibroblast growth factor (bFGF) have been investigated for applications in myocardial infarction (9,10), peripheral ischemia and wound repair (11-13). Platelet derived growth factor (PDGF) has also been studied for such uses and has been shown to recruit cell types important in the neovascularization process such as pericytes and smooth muscle cells (8). Stromal-cell derived factor-1 (SDF-1) has been used for similar applications and has demonstrated recruitment of bone-marrow-derived stem cells and endothelial progenitor cells (14,15). In addition, SDF-1, insulin-like growth factor-1 (IGF-1), and hepatocyte growth factor (HGF) are known to have angiogenic and cardioprotective effects (16-19). Many of these growth factors have affinity to binding sites on engineered scaffolds, therefore allowing for sustained release and enhanced efficacy at the site of application (20). Thus, growth factors are a powerful component of tissue engineering by enhancing vascularization and integration of tissue-engineered constructs as well as by promoting differentiation of host or implanted cells. Several reviews are available discussing this topic (21-23).
Host response to tissue engineering therapies
Injury after implantation of any foreign material leads to an inflammatory response in the host tissue. Tissue engineered constructs are no exception to this. In this section the cascade of events comprising the body’s innate immune response to injury will be discussed followed by specific immune responses to synthetic and naturally derived tissue engineered scaffolds as well as modifications for modulation of immune response.
Inflammation and wound healing
Inflammation is the body’s mechanism to contain, neutralize, dilute or wall off a foreign object (24). The cascade of events that occurs after injury or implantation is outlined in Fig. 3. The time course of the inflammatory process can be broken down into acute and chronic inflammatory response followed by granulation tissue and a potential fibrous encapsulation and/or foreign body reaction.
Acute inflammation occurs minutes to a few days after injury and is dependent on the extent of the damage. During this process leukocytes – predominantly neutrophils and also monocytes migrate into the injury site. The accumulation of these leukocytes occurs via a complex process involving the stimulation of leukocyte adhesion molecules as well as endothelial adhesion molecules or molecules that have a combined effects such as C5a, IL-1 and tumor necrosis factor (TNF), respectively. Neutrophils emigrate first using chemotactic mechanisms and are short-lived (24 – 48 hours), thus becoming the predominant cell type during that window. A day or two after insult the chemotactic factors for monocytes are predominant and monocyte emigration can continue for days to weeks. The infiltrated monocytes can differentiate into macrophages at the injury site and the number of macrophages has been shown to peak at one week but can remain at the site for months (25,26). The macrophage response to biomaterial implantation is complex. Recently the phenotypic and functional polarization of macrophages has been studied in more depth (27-30) and is similar to the Th1/Th2 polarization shown for lymphocytes (31). M1 macrophages and T1 lymphocytes are of a pro-inflammatory and cytotoxic phenotype and are associated with chronic inflammation. On the other hand, M2 macrophages and T2 lymphocytes are anti-inflammatory and encourage tissue repair, regeneration and beneficial remodeling (27). The neutrophils and macrophages main function is to phagocytose the foreign entity. However, in the case of biomaterial scaffolds, there is a large size disparity and a process called “frustrated phagocytosis” may occur. Though the leukocytes are unable to engulf the biomaterial, they do release enzymes to attempt degradation of the material, which may cause further damage to the tissue (32). Persistent inflammatory stimuli can lead to chronic inflammation, which may be due to the physical and chemical properties of the implant or due to motion in the implant site. This chronic inflammation stage mainly involves macrophages, monocytes and lymphocytes and is restricted to the site of implantation. If resolved, the process continues into granulation tissue, which is characteristic of a healing during the inflammation process and can be seen as early as 3-5 days after injury. This is a pink tissue with a soft granular appearance containing fibroblasts, which help in wound contraction. There are also new blood vessels due to neovascularization in the granulation tissue (24).
The overall aim of wound healing is to reestablish a homeostatic environment, which is termed as resolution. This resolution process can occur in many different ways, but does not occur with persistent chronic inflammation. With permanently implanted materials, resolution typically occurs by the process of encapsulation, where a fibrous capsule composed of spindle shaped fibroblasts and a small number of macrophages forms around the implant. A foreign body reaction can also occur with characteristic foreign body giant cells, which are multinucleated cells formed by the fusion of macrophages (24). While encapsulation can be considered a biocompatible response for many implants, for a tissue engineering scaffold, this would be deleterious as it would wall of the scaffold and prevent proper tissue ingrowth and/or regeneration. Though rare, integration is another method of resolution for a permanently implantable material in which there is approximation of the host tissue with the implant, without capsule formation. When in contact with epithelial linings such as the skin, resolution can take place by the process of extrusion in which a pouch is formed around the material and it is forced out of the tissue, which again would not be desired for tissue engineering. Alternatively, resorption can occur when materials degrade at a rate fast enough to make sure that no fibrous capsule is formed around the implant. In this scenario, the timely degradation of the material results in replacement by the appropriate tissue, which is the goal for tissue engineering.
Immune response to tissue engineered therapies and modifications for modulation of immune response
Tissue engineering aims to replace injured or diseased tissue by functional tissue. With biomaterials being a major tissue engineering component, host inflammatory response after biomaterial implantation is an important consideration that determines biocompatibility and cell survival. In addition to the biomaterial, delivery of cells or growth factors can also stimulate an immune reaction that can have both positive and negative effects on tissue repair, regeneration and growth (31). These tissue engineering components can also potentially elicit a foreign body reaction and introduce antigens. Prevention of macrophage infiltration – a major inflammatory cell type, has been shown to be detrimental to regeneration and tissue salvage (33). Therefore, modulation of immune response to harness the beneficial properties of inflammation while mitigating the deleterious effects should be an aim of tissue engineering strategies.
There are two main types of immune responses that can occur – innate immune response, that is non-specific and present from birth (i.e. inflammation) or an acquired immune response that is specific for a particular foreign agent known as an antigen. Acquired immune response can be further delineated into antibody mediated (humoral) immune response and cell-mediated immune response. Many tissue engineered scaffolds are from naturally derived sources. These natural materials are mostly ECM derived and may be of allogeneic or xenogeneic sources. They may contain proteins that serve as antigens, which can be recognized as foreign by host antibodies, and thus initiate an antibody-mediated response. Alternatively, synthetic materials may have degradation components that promote a cell mediated immune response (34). The degradation components can be digested and engulfed by macrophages and the macrophages can then present surface receptors that are recognized by activated cytotoxic T-cells or killer T-cells that destroy the macrophages containing the foreign entity. In addition to the biomaterial component, introduction of cells can also induce an immune response. Cells can be derived from a variety of allogeneic or xenogeneic sources and may be recognized as non-self, eliciting a cytotoxic T-cell mediated response. Use of a patient-specific cell type may be of benefit, however in most cases allogeneic cell sources are used, where the risk of rejection is increased.
Though immune response is a valid concern, there are many ways in which the response can be modulated. For example, the surface properties of a biomaterial can be tailored to present greater hydrophilicity, which has been shown to decrease monocyte and macrophage adhesion and limit foreign body giant cell formation (35,36). Recently, decellularized ECM based scaffolds have been shown to promote a more pro-remodeling, anti-inflammatory response with increased M2 macrophages and Th2 lymphocytes (28,37,38). Porosity of the biomaterial may also play a role in the immune response elicited. Materials with a high surface to volume ratio have a greater macrophage and foreign body giant cell response, while materials with a pore size of 30 – 40 µm have shown greater vascularity and less fibrotic encapsulation as well as a shift towards the M2 macrophage phenotype. When the pore size is larger or smaller than the range above, a classical foreign body response is seen regardless of the origin of the biomaterial (39-41). Microtemplating techniques have been used to create materials with a uniform pore size and have been successfully utilized in cardiac (41) and dermal (39) tissue engineering applications.
Scope and future of tissue engineering
Tissue engineering approaches are contributing to the therapies and products available for the treatment of various conditions or diseases ranging from congenital mal-formations to left ventricular remodeling after myocardial infarction. Between 1990’s and 2001 was the golden era for the early promise of tissue engineering with greater than 70 start-up companies and over $600 million in expenditure (6). Many products were developed to augment organ replacement strategies due to paucity of donor organs (7). The future of tissue engineering looked bright, however in 2001-2003, tissue engineering had many severe setbacks during the dot-com crash. This was partially due to growing skepticism of investors in conjunction with poor clinical trial outcomes and business plans. Some of the major tissue engineering companies of the time such as Advanced Tissue Sciences and Organogenesis were forces into bankruptcy (6). Nonetheless, tissue engineering made an impressive recovery in the last decade. Part of this can be accredited to the diversification in tissue engineering products, most notably the increased development and use of acellular products along with a better understanding of business models. Over the last couple years we have seen many of these products reach commercialization; in 2007 tissue engineering and regenerative medicine products generated profits greater than $1.3 billion that benefited about 1.2 million patients (6). With many products already on the market and even more in clinical trials, the success and continuity of tissue engineering products hinges on a thorough understanding of the mechanisms by which these therapies provide functional benefit.
C. Cardiac Tissue Engineering
With over 1 million Americans suffering from MI each year (42), novel therapies for the treatment of MI and subsequent heart failure is the need of the day. Current therapies include pharmacologics, left ventricular (LV) assist devices and total heart transplantation, however many experimental tissue engineering approaches are emerging for the treatment of adaptive LV remodeling and subsequent HF after MI. For approximately two decades, cellular cardiomyoplasty has been explored as a potential therapy, and there are numerous ongoing clinical trials with some Phase III studies just beginning. Cell injection has, however, been limited by very poor retention and survival. A large portion of the transplanted cells can leak out of the tissue due to low viscosity of the injection medium (culture media or saline) contributing to poor retention of the cells (9). In addition, cell survival is limited due to the hostile infarct environment, possible immune rejection by the host, and lack of a suitable matrix for cell adherence (43). Studies have shown that only about 10% of the injected cells are retained in the myocardium at 2 hours after injection (44,45) and a separate study has shown that 90% of the retained cells died due to the inhospitable environment (46). Initially, tissue engineering approaches were explored in the heart to combat these findings by providing an appropriate biomimetic environment for the injected cells to increase both engraftment and survival; moreover, more recently, acellular scaffold approaches that encourage endogenous cell infiltration are being utilized. Traditionally, an epicardial patch-based approach was utilized to deliver cells to the epicardial surface of the heart, which requires an open chest surgical procedure for implantation. To develop a potentially minimally invasive method for delivery, in situ gelling materials with or without cells or growth factors have been pursued, with some of these materials being catheter deliverable (47,48). With the positive results seen in basic science and animal pre-clinical research, and most recently clinical trials, the future for cardiac tissue engineering looks bright.
Both cellular and acellular biomaterial approaches for cardiac tissue engineering (Fig. 4) have shown improved cardiac function, reduced infarct scar and increased neovascularization in the infarction zone. With the promising results seen from such therapies, many have moved into large animals pre-clinical trials and we are beginning to see these therapies move toward human application. In the subsequent discussion, seminal findings pertaining to cellular and acellular patch and injectable approaches for cardiac tissue engineering are discussed. A summary of patch and injecatable cardiac technologies can be found in Table 1. More comprehensive reviews on in vivo studies discussing cardiac tissue engineering are available (9,10).
Cardiac tissue engineering consists of cardiac patches and injectable biomaterials. Cardiac patches and injectable materials can be either used as acellular scaffolds (A, D), or delivery vehicles for cells (B, E) and/or biological molecules (C, F). Reproduced from (10) with permission from Elsevier.
Table 1. A Review of Cardiac Tissue Engineering Therapies. Modified and reprinted with permission from (9) and (10).
Cardiac patches can be used as a vehicle for delivery of cells to the epicardial surface of the infarct and also to provide the microenvironment necessary to promote cell survival or alternatively, as an acellular therapy to provide the necessary structural support and/or the biological cues required to promote repair and/or regeneration. Many cardiac patch studies have shown an improvement in cardiac function and LV dimensions. Select studies demonstrating the use of cellular and acellular cardiac patches for the attenuation of LV remodeling and HF after MI are discussed.
Various biomaterial patches such as those derived from gelatin, alginate, collagen, Matrigel, and a few synthetic materials, with a variety of different cells types such as cardiomyocytes, dermal fibroblasts, bone marrow-derived mesenchymal progenitor cells, and embryonic stem cells have shown an improvement in cardiac function in small animals (10). The most progress in the field has been made with collagen-based scaffolds. Collagen patches have been seeded with a variety of cell types such as human bone marrow-derived CD133+ cells leading to increased angiogenesis (49) or human mesenchymal stem cells (MSCs) causing improved cardiac function (50). Patches using a combination of different biomaterials were also used in conjunction with cells. Neonatal cardiomyocytes and pro-survival and angiogenic factors in Matrigel™ medium were delivered using an alginate scaffold. The patch was cultured in the rat omentum to promote maturation of vasculature and then implanted in a rat MI model. This complex patch showed functional integration with the host myocardium through the formation of gap junctions (51).
A technique called decellularization has also been used for the creation of patches. Decellularization allows for the removal of the native tissues’ cellular content (52). Human mesenchymal progenitor cells in fibrin were seeded on a decellularized human myocardial sheet. Implantation of the scaffold over the infarct in a nude rat model demonstrated preservation of cardiac function as well as an enhanced vascular network (53). Similar results were seen with application of a patch derived from decellularized small intestinal submucosa (54) and decellularized pericardial tissue (55) seeded with MSCs.
Okano et al. utilized a temperature responsive synthetic polymer, poly(N-isopropylacrylamide) (PNIPAAM) for the novel application of creating cells sheets that can be lifted off a substrate and then used as a patch for cardiac repair. They created constructs of a thickness up to 600 µm by stacking multiple sheets of MSCs (56,57). Implantation in a rat MI model resulted in improved cardiac function and infarct thickness (56). Since then similar results have been obtained using a variety of different cells types such cardiac progenitor cells (58), human menstrual derived mesenchymal cells (59) and skeletal myoblasts (60) as well as a co-culture of neonatal cardiomyocytes and endothelial cells (61) for cell-sheet based technologies.
In the Myocardial Assistance by Grafting a New Bioartificial Upgraded Myocardium (MAGNUM) Phase I clinical trial, a collagen type I patch seeded with autologous bone marrow cells was applied on 10 patients that had coronary artery bypass graft and stem cell therapy. This combined approach of cell injection and a cardiac patch resulted in an improvement in New York Heart association functional class, wall thickness and cardiac function compared to cellular injection alone (62). Thus indicating that this combined approach of delivering cells through injection and a patch or cell sheet may be beneficial.
More recently, there has been an emergence of cardiac patches as acellular therapies in addition to vehicles for growth factor delivery. Collagen patches alone (63) or with growth factors such as VEGF (64) have also been used for acellular treatments in a rat cryoinjury model, displaying increased angiogenesis compared to the control infarct. In addition to biologically derived materials that may contain the biological factors necessary to promote a cellular response, synthetic materials have also been assessed as potential scaffold materials. In a study by Robinson et al., a urinary bladder derived ECM patch or an expanded polytetrafluoroethylene patch was applied to the epicardial surface of a porcine infarct model. Three months after application, the ECM patch was completely reabsorbed and demonstrated greater cell migration than the non-degradable synthetic scaffold that elicited a foreign body response with necrosis and calcification (65). On the other hand, Fujimoto et al. showed that application of a biodegradable elastic polyester urethane urea cardiac patch over an infarct led to improvement in fractional area change and wall thickness at 8 weeks as well as an increase in muscle bundles in the infarct area. The patch degraded over the course of 8 weeks and likely served as a biocompatible structural support allowing mechanical augmentation of the damaged wall while at the same time avoiding potential complications such as chronic inflammation that can occur with a non-degradable material (66).
Considerations for cardiac patch technology
In order to generate tissue engineering patches that are likely to improve contractile function in the human heart, a myocardial thickness on the order of 1 cm is necessary. However, diffusion limitations are a significant challenge in the development of such tissue constructs. Currently, tissue constructs of 600 µm thickness have been generated; while this may be applicable in a rodent model this may serve as a hurdle to human application. However, in a recent pre-clinical trial Miyagawa et al. showed that application of a 200 µm – 300 µm skeletal myoblast cell sheet to a porcine infarction led to increased systolic and diastolic function and myocardial perfusion, reduced fibrosis and cell diameter, plus an increase in vessel density (67); indicating that perhaps even a thin patch may lead to improvement in larger hearts, likely through secretion of paracrine factors. In addition, for the use of patches in cellular delivery, vascularization of the patch is critical to ensure survival of the patch post-implantation. However, this remains a challenge in many cases. Cardiac tissue has a highly organized structure with changing fiber directions though the heart. An ideal patch will correctly mimic the healthy myocardial fiber angles and direction. To date, the current biomaterials have yet to provide a mimic to the complex fiber orientation in the heart, which may be important for cardiac force production during systole as well as electrical integration with the host tissue.
To overcome some of the shortcomings seen with the cardiac patch approach, biomaterials have been injected directly into the vulnerable areas of the myocardial wall. These approaches have the added benefit of being potentially minimally invasive along with being able to target the entire infarction region instead of just the epicardial surface as seen with the patch-based approach. A summary of the work in the field of injectable biomaterials is provided below.
Injectable biomaterials were first used as a vehicle for cell delivery for MI therapy. Fibrin was the first injectable biomaterial to be tested for such an application. In a study by Christman et al, skeletal myoblasts in a fibrin scaffold were delivered via an intramyocardial injection into an occlusion-reperfusion rat MI model. Injection of cells in the biomaterial scaffold improved cell transplant survival, reduced infarct size and led to increased neovascularization (68) as well as preservation of cardiac function and wall thickness (69). In addition, improvement in cardiac function and cellular retention has been seen with many other cell types such as marrow-derived cardiac stem cells (70) and adipose-derived stem cells (71,72) injected with fibrin. Surprisingly, functional benefit was seen in the groups injected with fibrin alone as well as cells in fibrin, paving the way for acellular treatments. One potential drawback, similar to many other currently used injectable materials, is that it gels rapidly, rendering this an invasive therapy. Along with fibrin, materials like Matrigel have been injected with embryonic stem cell derived cardiomyocytes and pro-survival factors to in a rodent MI. This biomaterial/cell delivery system resulted in not only an increase in cell survival but also an improvement in ventricular geometry and cardiac function (73).
Bio-inspired materials have also been used as a method for cell delivery (74-76). Work by Davis et al. demonstrated that biotinylated peptide nanofibers tethered to IGF-1 together with neonatal cardiomyocytes improved systolic function in a rat MI model (74). Since then, other studies have followed suit and shown similar results in larger animals (76). Most recently, synthetic materials have been utilized for these purposes as they have the added benefit of control over properties such as degradation, stiffness, porosity and gelation times as well as limited batch-to-batch variability that can be seen in materials derived from a biological source. For example, implantation of a matrix metalloproteinase degradable Poly(NIPAAM-co-acrylic acid) with bone marrow – derived mononuclear cells led to an improvement in function; however the injection of the scaffold alone caused greater functional improvement than the scaffold with cells. The timely degradation and cellular infiltration may have contributed to this effect.
Many studies utilizing a cell and biomaterial construct have also included a biomaterial only group and have shown functional benefit without the presence of cells. This has paved the way for the plethora of acellular injectable treatments for MI, some of which are outlined below and others are discussed in Table 1.
Many different biomaterial scaffolds have been used for the treatment of post-MI remodeling, however the greatest progress in terms of clinical translation has been made with alginate – a seaweed derived polysaccharide. Christman et al, laid the groundwork for injectable biomaterials as a treatment for MI with fibrin, however fibrin’s rapid gelation makes minimally invasive delivery a challenge (10,47) given that the catheter would either clog or have to be repeatedly flushed. On the other hand, a catheter – safe delivery method has been established with Alginate. The work done by Landa et al., showed that injection of alginate in acute and chronic rodent infarcts leads to increased scar thickness and mitigated dilatation and dysfunction (77). Incorporation of growth factors such as VEGF (78), IGF and HGF (19) as well as modification with peptides such as RGD (79-81) has demonstrated a beneficial effect on the LV remodeling cascade. With alginate showing positive results in small animal models, pre-clinical work in a swine model was carried out showing that intracoronary injection of alginate led to prevention and reversal of negative LV remodeling (48). Most recently, alginate underwent a Phase I safety and efficacy trial (82), and will be further examined in the Phase II PRESERVATION I (IK-500 for the Prevention of Remodeling of the Ventricle and Congestive Heart Failure After Acute Myocardial Infarction) trial.
As seen in the development of cardiac patches, decellularization is also being used as a technique for injectable biomaterial development. These materials provide the added benefit of carrying the biochemical cues of the native ECM. Injection of decellularized small intestinal submucosa extracellular matrix into a rodent MI promoted cell migration and improvement in ejection fraction (83). The organ from which the decellularized material is made may be of importance, as the ECM is known to have a heterogeneous composition of proteins and proteoglycans (84,85). Hence, a biomaterial derived from the myocardium may contain the necessary composition for cardiac tissue engineering. Singelyn et al. developed a liquid form of decellularized porcine ventricular tissue that self-assembles into a gel upon injection. Injection of this material in a MI showed preservation of cardiac function in rats without adversely affecting electrical signal propagation. This material has also been shown to be compatible with transendocardial catheter delivery (47) and positive results were recently shown in a porcine model of MI (unpublished), demonstrating its potential clinical applicability.
Another biomaterial, hyaluronic acid was tailored to different stiffness and injected into an ovine infarct. Though there were no significant differences in cardiac functional parameters, injection of the stiffer material demonstrated reduction in infarct area (86), thus identifying material mechanics as a potentially important factor. Computational modeling techniques have also been utilized to better understand the impact of material mechanics on cardiac function after MI and suggested that presence of a material could increase wall thickness and hence decrease wall stress due to Laplace’s Law, thus mitigating negative LV remodeling and contributing to improvement in cardiac function (87).
Compared to naturally derived materials, synthetic polymers have an added benefit of being able to decouple bioactivity and material mechanics as well as allowing for controlled delivery of biological moieties for tissue engineering applications. Rane et al. showed that injection of a bio-inert, non-degradable material, did in fact increase wall thickness; however, it was insufficient to prevent the deleterious effects of LV remodeling (88). This suggests that inherent bioactivity of the material or timely degradation of the material allowing for cell infiltration may be potential mechanisms by which injectable scaffolds affect cardiac function. Synthetic materials have also been explored as scaffolds for delivery of growth factors such as basic fibroblast growth factor (89), VEGF (90) and erythropoietin (91).
Considerations for injectable biomaterial technology
As mentioned previously, translatability of injectable biomaterial therapies largely depends on their ability to be delivered minimally invasively via a catheter. Most injectable materials have been studied in the rodent heart. However, translation of this therapy to large animals or humans requires either multiple injections in the infarcted myocardium or infusion through the damaged vasculature. Many injectable materials gel at a rate that makes multiple injections via a catheter, prior to clogging, unfeasible. Others would not be hemocompatible for vascular infusion. Therefore, new injectable biomaterials should be designed to allow for catheter delivery. The selection of which therapy is most applicable may be dependent on infarct characteristics, such as size and transmurality of infarction and duration post-infarction. Chronic infarcts are replaced by collagen deposition, causing thinning of the wall and an increase in stiffness of the tissue. Tissue engineered cardiac patches may provide therapeutic benefit by supporting the already weakened wall by application to the epicardial surface or removing the scar tissue by surgical ventriculotomy (92).
Design parameters and future considerations for biomaterials therapies
Cellularity and bioactivity of the scaffold
Both biological and synthetic biomaterials have been utilized for cellular or acellular cardiac applications. However, it is still unclear as to whether addition of cells causes functional improvement over the acellular scaffold. The results are mixed and perhaps this can be attributed to differences in experimental conditions. For example, current literature shows both bone marrow mononuclear cells seeded on a degradable poly-glycolide-co-caprolactone scaffold and the acellular scaffold reduced LV dilatation and preserved LV systolic function (93). Similar results were seen with other scaffolds as well. Opposite to these findings, application of a polylactide-co-€- caprolactone patch alone was unable to improve EF or decrease infarct size (94). However, the majority of injectable therapies have shown similar improvements in function and remodeling regardless of the cellularity of the construct. As highlighted in the aforementioned sections of this chapter, both synthetic and biological materials have been tested as potential treatments for post-MI LV remodeling, with both materials showing promising outcomes. Biologically derived materials have the benefit of providing the adequate biochemical environment necessary for cellular recruitment to the damaged area and may better mimic the heart’s extracellular environment. Then again, synthetic materials have the added benefit of material tunability to match complex cardiac mechanical properties and physical structure, such as pores.
Structural characteristics of the scaffold
The purpose of cardiac tissue engineering strategies for MI treatment is to mitigate tissue damage caused during infarction as well as curb processes such as infarct expansion and further deterioration of the injured myocardium. During the LV remodeling cascade there are structural changes in the LV such as wall thinning and collagen deposition leading to changes in stiffness of the infarct. With the different structural and mechanical changes in the infarct, stiffness of the biomaterial implant is an important consideration. In the case of injectable biomaterials, materials ranging from Pa range to the kPa range have been developed and tested. Berry et al. have shown the elastic modulus of the healthy myocardium is around 20 kPa (95), so it might be beneficial for biomaterials to have mechanical properties in this regime. On the other hand, individual ECM components lie in the 0.5 to 1 kPa range and may also have advantageous effects. Ifkovits et al. studied the effect of changing material mechanics on cardiac function after MI. They demonstrated that injection of a stiffer material (43 kPa) close to the stiffness of infarcted tissue (55 kPa) (95), reduced scar size, but did not affect cardiac function (86). Though ECM derived biomaterials have lower mechanical properties, techniques such as chemical crosslinking can be used to enhance the mechanical properties of the scaffold (96). However, further studies investigating the effect of material mechanics on LV remodeling could provide some answers on how to best modulate mechanical properties of biomaterial treatments.
Another important characteristic of biomaterials is degradation rate. All of the bioactive and majority of the synthetic biomaterials used for cardiac tissue engineering applications are degradable in the range of 1 to 6 weeks (10). Apt degradation of biomaterials may allow for timely cell infiltration allowing for repair of the infarct and salvage of the vulnerable borderzone area. These infiltrating cells may be valuable for increasing neovascularization and modulating the inflammatory response, with perhaps a greater M2 macrophage and T2 leukocyte response. In addition, the timely degradation may allow for transient changes in mechanical properties allowing for the stiffness to change as the material degrades. This may be beneficial in the remodeling infarct by initially providing support to the vulnerable tissue but then degrading over time to allow for cellular infiltration.
Time point of delivery and assessment
Currently, there is heterogeneity in time points for intervention with cardiac tissue engineering strategies. Intervention with injectable biomaterials in preclinical models is most commonly carried out immediately and 1 week after infarction. Immediately after infarction the myocardial wall is susceptible to rupture and hence this time point may not be clinically relevant. Most patches are implanted 2 to 4 weeks after the infarction (10), which is in the middle of the rat remodeling cascade that is largely complete by 5 weeks. The time point for delivery is an important criterion prior to human translation. As of today, it is still unclear as to whether these therapies should be used for acute or chronic MI patients. In most of the studies carried out to date, tissue engineering therapies have been applied to acute or subacute infarcts and have shown beneficial effects. However, limited studies have applied tissue engineering approaches to the chronic infarct and shown improvement in cardiac function (77). Lately, there has been a greater push for reperfusion of MI patients, perhaps the most suitable time for delivery should mimic the time course of reperfusion. On the same token the impact of cardiac tissue engineering approaches have been assessed at various durations ranging with majority of studies evaluating effects at 4 weeks after biomaterial delivery. To better understand the efficacy of these approaches more studies at longer time points are needed. For example, injection of a synthetic polymer was shown to improve cardiac function at an early time point of 4 weeks but at later time point of 13 weeks that effect diminished (97).
Future of cardiac tissue engineering
The field of cardiac tissue engineering is advancing at a rapid pace. In the last 5 years both cardiac patch and injectable biomaterial treatments have reached the clinical trial stage and are showing promising outcomes in initial Phase I safety and feasibility trials. With the advent of in situ gelling biomaterials, allowing for minimally invasive modes of delivery, the accessibility to this emerging treatment should increase. As we move to human application it is exceedingly important that we understand the mechanism by which these materials influence LV remodeling, cardiac function, and how best to tailor future therapies to maximize their efficacy.
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