Chapter 40

Hepatic tissue engineeringAmanda X. Chen1,2,*, Arnav Chhabra2,3,4,*, Heather E. Fleming2,3 and Sangeeta N. Bhatia2,3,4,5,6,7

1Department of Biological Engineering, Massachusetts Institute of Technology, Cambridge, MA, United States, 2David H. Koch Institute for

Integrative Cancer Research, Massachusetts Institute of Technology, Cambridge, MA, United States, 3Harvard-MIT Health Sciences and Technology,

Massachusetts Institute of Technology, Cambridge, MA, United States, 4Institute for Medical Engineering and Science, Massachusetts Institute of

Technology, Cambridge, MA, United States, 5Wyss Institute for Biologically Inspired Engineering, Boston, MA, United States, 6Department of

Electrical Engineering and Computer Science, Massachusetts Institute of Technology, Cambridge, MA, United States, 7Howard Hughes Medical

Institute, Chevy Chase, MD, United States

Liver disease burden

Fatal liver disease accounts for B2 million deaths annu-

ally worldwide and has steadily increasing rates over the

years [1]. Liver failure can be divided into three major

categories: (1) acute liver failure (ALF) that presents as a

rapid loss of liver function in patients without preexisting

liver disease, (2) chronic liver disease due to metabolic

dysfunction, and (3) chronic liver failure accompanied by

tissue remodeling and scarring.

ALF is a rare syndrome with an annual incidence of

less than 10 cases per million people in the developed

world. In the United States, B2000 cases of ALF are

diagnosed each year [2]. It commonly develops in healthy

adults in their 30s. Patients with ALF usually present with

abnormal liver biochemistry, coagulopathy, and encepha-

lopathy. The causes vary geographically. Damage due to

drug exposure (e.g., acetaminophen) is the most common

cause in the West, while in large parts of the East, viruses

(e.g., Hepatitis A and E) are the most prominent cause of

ALF [3]. Clinically, ALF can be subdivided based on the

period of time between the appearance of jaundice and

onset of hepatic encephalopathy. Data from O’Grady

et al. proposed the following classification: hyperacute for

periods between 0 and 7 days, acute for periods between

7 and 28 days, and subacute for periods between 4 and 12

weeks [4]. In hyperacute cases the cause is usually acet-

aminophen toxicity or viral infection. Subacute cases that

evolve slowly often result from idiosyncratic drug-

induced liver injury (DILI). Even though patients with a

subacute presentation have less coagulopathy and enceph-

alopathy, paradoxically they have a consistently worse

medical outcome than those with a more rapid onset of

the disease [5].

Chronic liver disease develops on the background of a

constant injurious insult, either resulting from a metabolic

disorder or a number of etiologies that lead to widespread

tissue remodeling and pathologic deposition of extracellu-

lar matrix (ECM). Inborn liver-based errors of metabolism

are life-threatening conditions caused by genetic defects

in single enzymes or transporters and lead to blockade of

a specific metabolic pathway. While they can be accom-

panied by progressive fibrosis and cirrhosis, such as in

the case of α1-antitrypsin ZZ deficiency, hemochromato-

sis, Wilson’s disease, and hereditary tyrosinemia [6], the

liver parenchyma often remains intact. Some examples of

metabolic disorders with an intact parenchyma include

hypercholesterolemia, Crigler�Najjar syndrome, ornithine

transcarbamylase deficiency, organic acidurias, and

hyperoxaluria [7]. In a biopsy the lack of parenchymal

destruction often leads to a delayed diagnosis, exposing

the patient to sequelae. In all cases of liver disease the

lack of FDA-approved noninvasive biomarkers makes it

challenging to diagnose and treat liver diseases.

Chronic liver disease occurs in the setting of nonalco-

holic fatty liver disease (NAFLD). NAFLD is marked by

hepatic steatosis and is related to the presence of meta-

bolic syndrome in association with obesity, diabetes, and/

or arterial hypertension [8]. A subset of NAFLD patients

* These authors contributed equally.

737Principles of Tissue Engineering. DOI: https://doi.org/10.1016/B978-0-12-818422-6.00041-1

Copyright © 2020 Elsevier Inc. All rights reserved.

will develop signs of nonalcoholic steatohepatitis

(NASH), a more severe condition associated with lobular

inflammation and hepatocellular ballooning and can lead

to fibrosis and cirrhosis [9]. In NAFLD the liver is unable

to utilize carbohydrates and fatty acids properly, leading

to toxic overaccumulation of lipid species. These metabo-

lites induce cellular stress, injury, and death, which pre-

dispose the liver to sequelae such as cirrhosis and

hepatocellular carcinoma [10].

In the United States the number of NAFLD cases is

projected to expand from 83.1 million in 2015 (B26% of

the population) to 100.9 million by 2030 (B28% of the

population) [11]. An increasing percentage of these cases

are projected to be classified as NASH, rising from 20%

to 27% of adults with NAFLD during this interval [12].

While diagnosing NASH at an early stage remains a chal-

lenge, multiplexed protease-activated nanosensors have

demonstrated utility in monitoring NASH progression and

treatment response in a 3,5-diethylcarbonyl-1,4-dihydro-

collidine model of fibrosis in mice. With further develop-

ment, these noninvasive readouts can be used to diagnose

disease.

Current state of liver therapies

In order to mitigate the clinical burden of liver disease,

several therapeutic strategies have been undertaken

(Fig. 40.1).

Extracorporeal liver support devices

Liver failure is associated with abnormal accumulation of

numerous endogenous substances such as bilirubin,

ammonia, free fatty acids, and proinflammatory cytokines

[13]. Extracorporeal liver support devices have been

developed to detoxify the blood and plasma in order to

bridge patients to liver transplantation (LT) or allow the

native liver to recover from injury. Artificial liver (AL)

devices use nonliving components for detoxification, such

as membrane separation or sorbents, to selectively remove

toxins but have limited clinical use because they do not

replace the synthetic and metabolic roles of the liver [13].

Bioartificial liver (BAL) devices, on the other hand, con-

tain a cell-housing bioreactor that aims to provide the

detoxification and synthetic functions of the liver and are

an ongoing topic of clinical investigation. Current ver-

sions are either based on hollow fiber cartridges [14�16]

or on perfused three-dimensional (3D) matrices [17].

BALs, just like other hepatocyte-based therapies, face

many challenges, such as the lack of readily available

functional cell sources and the loss of cell viability and

phenotype during the treatment process.

Biopharmaceuticals

In the setting of ALF, N-acetyl cysteine (NAC) is FDA-

approved to reduce the extent of liver injury after acet-

aminophen overdose [18]. In the setting of chronic liver

disease, however, most of the FDA-approved therapies

are for hepatitis A, B, and C. A detailed listing can be

found in Table 40.1. While a few treatments have shown

moderate efficacy, there are currently no biopharmaceuti-

cals that are approved for NAFLD, NASH, or cirrhosis.

Glitazones, for example, upregulate adiponectin, an adi-

pokine with antisteatogenic and insulin-sensitizing proper-

ties [19]. Vitamin E, an antioxidant, can prevent liver

injury by blocking apoptotic pathways and protecting

against oxidative stress [19]. Despite clinical studies of a

large number of therapeutic candidates, no single agent or

combination has shown improvement to liver-related mor-

bidity and mortality in patients with NASH. Until a drug

is FDA-approved for NASH indications, lifestyle modifi-

cations and optimizing metabolic risk factors are the best

medical-treatment options for these patients.

Liver transplantation

The first attempt at human LT took place at the University

of Colorado on March 1, 1963 but turned out to be unsuc-

cessful [20]. Based on the pioneering work of Thomas

Starzl, the first extended survival of a human recipient after

LT was achieved on July 23, 1967 with a 19-month-old

female patient with hepatocellular carcinoma who survived

FIGURE 40.1 Cell-based therapies for liver disease. Extracorporeal

devices perfuse patient’s blood or plasma through bioreactors.

Hepatocytes are transplanted directly or implanted onto scaffolds.

Genetically modified large animals can be used for xenotranplantation.

738 PART | ELEVEN Gastrointestinal system

13 months before succumbing to metastatic disease [21].

After the initial success of the surgery, advancements were

made to improve donor organ quality, recipient selection,

operative and perioperative management, immunosuppres-

sion and infectious complications. These advancements have

made orthotopic LT the primary treatment for end-stage

liver disease and certain cancers. These transplants have

1-year patient survival rates over 80% [22]. However, many

challenges remain, including donor organ shortages,

recipients with more advanced disease at transplant, a grow-

ing need for retransplantation, and adverse effects associated

with long-term immunosuppression. To overcome a growing

imbalance between the supply and demand of donor livers,

transplant centers have developed strategies to expand the

donor pool. These strategies include live donor LT [23],

split-LT [24], and extended criteria for donor livers [25].

Despite all these efforts, the number of liver transplants has

not increased in the last decade.

TABLE 40.1 List of FDA-approved therapies for chronic liver diseases.

Drug

name

Year

approved

Indication(s) Mechanism

Heplisav-B 2017 Hepatitis B Combines hepatitis B surface antigen with a proprietary Toll-likereceptor 9 agonist to enhance the immune response

Mavyret 2017 HCV genotype 1�6 Fixed-dose combination of glecaprevir, an HCV NS3/4A proteaseinhibitor, and pibrentasvir, an HCV NS5A inhibitor

Vosevi 2017 Hepatitis C Fixed-dose combination of sofosbuvir, an HCV nucleotide analogNS5B polymerase inhibitor, velpatasvir, an HCV NS5A inhibitor, andvoxilaprevir, an HCV NS3/4A protease inhibitor

Ocaliva 2016 Primary biliary cholangitis FXR agonist

Zepatier 2016 HCV genotype 1 or 4 Fixed-dose combination product containing elbasvir, an HCV NS5Ainhibitor, and grazoprevir, an HCV NS3/4A protease inhibitor

Cholbam 2015 Bile acid synthesis andperoxisomal disorders

Primary bile acid synthesized from cholesterol in the liver

Daklinza 2015 HCV genotype 3 Inhibitor of NS5A, a nonstructural protein encoded by HCV

Technivie 2015 HCV genotype 4 Fixed-dose combination of ombitasvir, an HCV NS5A inhibitor,paritaprevir, an HCV NS3/4A protease inhibitor, and ritonavir, aCYP3A inhibitor

Olysio 2013 Hepatitis C Small molecule orally active inhibitor of the NS3/4A protease of HCV

Sovaldi 2013 Hepatitis C Inhibitor of the HCV NS5B RNA-dependent RNA polymerase

Incivek 2011 HCV genotype 1 Inhibitor of the HCV NS3/4A serine protease

Victrelis 2011 HCV genotype 1 Inhibitor of the HCV NS3 serine protease

Viread 2008 Hepatitis B Oral nucleotide analogue DNA polymerase inhibitor

Tyzeka 2006 Hepatitis B Inhibitor of HBV DNA polymerase

Baraclude 2005 Chronic hepatitis B withevidence of active viralreplication

Small-molecule guanosine nucleoside analog with selective activityagainst HBV polymerase

Hepsera 2002 Chronic hepatitis B withevidence of active viralreplication

Inhibitor of HBV DNA polymerase

Pegasys 2002 Chronic hepatitis C withcompensated liver disease

Binds to and activates human type 1 interferon receptors

Peg-intron 2001 Chronic hepatitis C Binds to and activates human type 1 interferon receptors

Ribavirin 2001 Chronic hepatitis C Synthetic nucleoside analog with antiviral activity

Twinrix 2001 Hepatitis A and B Recombinant vaccine

FXR, Farnesoid X receptor; HBV, hepatitis B virus; HCV, hepatitis C virus; NS3, nonstructural protein 3.

Hepatic tissue engineering Chapter | 40 739

An alternative to human LT is xenotransplantation,

though it has been clinically intractable due to concerns

about immunological rejection and zoonotic pathogen

transfer. With the advent of accessible genetic engineer-

ing technologies to circumvent the aforementioned chal-

lenges, the breeding efficiency of animals can be

leveraged to mass-produce tissue for human organ trans-

plants. Niu et al. applied CRISPR-Cas9 to inactivate all

62 copies of porcine endogenous retroviruses, thus paving

the way for pig-to-human transplants [26]. Relatedly,

Langin et al. genetically engineered porcine heart xeno-

grafts and demonstrated long-term pig-to-baboon orthoto-

pic transplantation [27].

Hepatocyte transplantation

Given the several drawbacks of LTs, alternative strategies

have been pursued. A potential alternative to LT is alloge-

neic hepatocyte transplantation (HT). Transplanted cells

can provide the missing or impaired hepatic function once

engrafted. Given their synthetic and metabolic capabili-

ties, mature hepatocytes are the primary candidates for

liver cell transplantations. HT offers several advantages

over LT. It is less invasive and can be performed repeat-

edly to meet metabolic requirements. Furthermore, multi-

ple patients can be treated with a single dissociated donor

tissue, and harvested cells can be cryopreserved for later

use on an as-needed basis.

The first experimental attempt of HT was done in

1976 to treat an animal model for Crigler�Najjar syn-

drome type I [28]. Along with other observations, it led to

the first transplant of autologous hepatocytes in 10

patients with liver cirrhosis in 1992 in Japan [29]. Since

then, reports have been published on more than 100

patients with liver disease treated by HT worldwide [30].

Human HT has resulted in partial correction of a number

of liver diseases, including urea cycle disorders [31], fac-

tor VII deficiency [32], glycogen storage disease type 1

[33], infantile Refsum disease [34], phenylketonuria [35],

severe infantile oxalosis [36], and ALF [37]. HT faces

several limitations: limited supply of high-quality mature

hepatocytes, freeze�thaw damage due to cryopreserva-

tion, poor cellular engraftment (estimated to be from

0.1% to 0.3% of host liver mass in mice after infusion of

3%2 5% of the total recipient liver cells) [30], and allo-

geneic rejection.

Clinically, the most widely used administration route

for HT is through the portal vein or one of its branches.

Hepatocytes traverse the sinusoidal vasculature and create

transient occlusions. The occlusions lead to vascular per-

meabilization which allows transplanted cells to reach the

liver parenchyma [35]. The number of cells that are

injected intraportally and subsequently engraft is a func-

tion of portal pressure and liver architecture. Thus other

administration routes have been explored for patients with

cirrhosis who have high portal pressures due to fibrosis.

In animal studies, hepatocytes transplanted into the

spleen proliferate for extended periods of time and display

normal hepatic function. The spleen has been shown to be

well-suited for hepatocyte engraftment because it func-

tions as a vascular filter and provides an immediate blood

supply [30]. The peritoneal cavity represents an attractive

administration route as it is easily accessible and can

house a large number of cells. Due to cell number

requirements associated with metabolic compensation, it

has been used in patients with ALF [38]. As an alternative

to the portal vein, spleen and peritoneum, the lymph node

(LN) has also been shown to demonstrate engraftment of

donor hepatocytes [39,40]. While this strategy has not

been utilized in the clinic yet, the preclinical data is

promising.

Current clinical trials

Several pathways have been implicated in the biology and

pathogenesis of NAFLD development: insulin resistance,

lipotoxicity, oxidative stress, altered immune/cytokine/

mitochondrial functioning, and apoptosis. New therapeu-

tic modalities are being developed to target many of these

pathways. For a detailed overview of NAFLD-targeted

drugs that are currently in the clinical trial pipeline, please

refer to Younossi et al. [41].

In vitro models

To build high-fidelity cellular models and therapies, com-

ponents of the native liver microenvironment must be

incorporated (Fig. 40.2). The liver’s highly organized

structure is key to its role as a complex tissue supporting

myriad synthetic and metabolic functions. In addition to

hepatocytes the main parenchyma of the liver, there are

several nonparenchymal cell types such as liver sinusoidal

endothelial cells, Kupffer cells, cholangiocytes, and stel-

late cells. In each lobule of the liver an array of parallel

hepatocyte cords are sandwiched between the sinusoid,

carrying circulating blood, and the bile duct, carrying

hepatocyte-secreted bile acids. Notably, this arrangement

dictates a unique set of architecturally driven cell�cell

and cell�matrix cues, which gives rise to liver-specific

phenotypes. Gradients of physicochemical stimuli along

the sinusoid drive zonal phenotypes with disparate meta-

bolic and synthetic functional profiles [42]. Interrupting

the natural order of cell arrangement in the liver is

directly connected with diseases discussed in the “Liver

disease burden” section. In this chapter, we will primarily

focus on human platforms, which are biologically distinct

from animal-derived cell models of the liver that are

reviewed more in depth elsewhere [43].

740 PART | ELEVEN Gastrointestinal system

Two-dimensional liver culture

Hepatocytes are responsible for more than 500 metabolic

and synthetic functions of the human body, often catego-

rized broadly as protein synthesis and secretion, detoxifi-

cation, bile synthesis, and nitrogen metabolism. Primary

hepatocytes quickly lose their phenotype and function

after a few days in traditional monolayer culture and

require a collagen-coated surface for adherence and sur-

vival [6]. In contrast, when primary human hepatocytes

(PHHs) are cultured between two layers of collagen gel

(i.e., sandwich culture configuration), they retain viability,

polarity and many axes of relevant metabolic and syn-

thetic function [44,45]. Guguen-Guillouzo et al. found

that a random coculture with a liver epithelial cell line

was sufficient to support hepatic albumin secretion, sug-

gesting the importance of heterotypic cell interactions for

long-term ex vivo culture [46]. Bale et al. demonstrated

that other nonparenchymal liver cells can better recapitu-

late hepatic response to inflammatory stimuli, through

higher order intercellular interactions captured in a multi-

cellular platform [47]. It was later discovered that a

micropatterned architecture consisting of hepatocyte-filled

islands surrounded by a nonbiomimetic cell type and

mouse fibroblasts can also stabilize hepatocytes, suggest-

ing the existence of conserved coculture signals across

species. These micropatterned cocultures (MPCCs) enable

the study of DILI and hepatotropic pathogen infection for

several weeks in vitro [48�50]. Furthermore, Davidson

et al. added hepatic stellate cells to the traditional MPCC

to create an in vitro model of NASH [51].

Despite the utility of two-dimensional (2D) hepatic

cultures in screening assays, a wealth of literature sug-

gests that they are dissimilar to hepatocytes in vivo.

Specifically, 2D formats, even with overlaid collagen

matrix, are more flattened than their native cuboidal

architecture. On a subcellular level, this translates to

major differences in cytoarchitecture, which is linked to

aberrant polarization and nonphysiological behavior

[52,53]. Griffith and Swartz have described the improved

presentation of relevant biochemical and mechanical cues

in 3D cultures, typically cell-laden hydrogels, compared

to traditional 2D cultures [54].

Three-dimensional liver constructs

Commonly, 3D hepatic cultures consist of primary cell or

induced pluripotent stem cell (iPSC)-derived spheroid and

organoid cultures, which are typically embedded in ECM-

based hydrogels [55]. Spheroids and organoids can be

manufactured using a variety of techniques, such as

microwell mold-based technologies, which offer a high

degree of composition and size control but are difficult to

scale. Spinning flasks and bioreactors can produce large

populations of spheroids, though they are typically non-

uniform in size and function. Bell et al. showed that pri-

mary human hepatic spheroids fabricated and cultured in

microwell plates serve as useful models of hepatotoxicity

and multiple liver pathologies [56].

Toward transplantable cell therapies, Stevens et al.

used microwell molds to create 3D hepatic spheroid

FIGURE 40.2 Advances in hepatic tissue engi-

neering. Traditional tissue culture approaches such as

the addition of extracellular matrix, soluble factors,

cocultivation with supporting cell types, hanging

drop, microwell molding, and nonadhesive surfaces

have enabled the early study of hepatocyte phenotype

in vitro in both 2D and 3D cultures. The advent of

technologies from disciplines such as chemical engi-

neering and electrical engineering has led to a new

level of control for hepatic tissue cultures, such as

micropatterning to template cell interactions, micro-

physiological systems to study the impact of bioac-

tive perfusate, polymeric biomaterials for

constructing 3D cell-laden grafts, perfusion technolo-

gies for decellularization/recellularization strategies,

and 3D printing for scalable engineering of cellular

grafts. 2D, Two-dimensional; 3D, three-dimensional.

Hepatic tissue engineering Chapter | 40 741

cocultures of PHHs and fibroblasts, which can be embed-

ded in agarose, fibrin, or polyethylene glycol hydrogel

scaffolds. The resulting tissue constructs support hepatic

function in vitro and in vivo after ectopic transplantation

into the peritoneal cavity [57]. Furthermore, Stevens et al.

demonstrated that implanting a tissue seed, consisting

of hepatic aggregates and vascular cords, into an

FRNG [fumarylacetoacetate hydrolase-deficient (Fah2/2),

recombinase activating gene-deficient (Rag12/2), nonob-

ese diabetes (NOD), and interleukin-2 receptor γchain-deficient (Il2rγ-null)] mouse model of hereditary

tyrosinemia leads to a 50-fold expansion in serum human

albumin and formation of perfusable vessels after 80 days

of cycled exposure to regenerative stimuli [58]. Relatedly,

Takebe et al. constructed liver buds consisting of iPSC-

derived hepatocyte-like cells (HLCs), mesenchymal stem

cells, and human umbilical vein endothelial cells; mesen-

teric transplantation of these human liver buds resulted in

vascularization and rescued a lethal TK-NOG mouse

model of liver injury [59].

To create clinically viable engineered liver constructs,

cell sourcing (discussed further in the “Cell sourcing” sec-

tion), and clinical-scale manufacturing of organoids and

spheroids must be addressed. Toward this end, Takebe

et al. constructed a large-scale liver bud microwell culture

platform, enabling the formation of 108 liver buds [60].

Physiological microfluidic models of liver

Despite improvements to in vitro liver platforms as model

systems and implants, static cultures lack physiologically

relevant dynamic components. The native liver’s dynamic

physiology arises from blood circulation and multiorgan

cross talk. Thus to improve biological fidelity, many have

attempted to create microfluidic models of the liver [61]. By

leveraging techniques from the semiconductor manufactur-

ing industry such as soft lithography, groups have fabricated

microphysiological systems with preformed channels to

allow for perfusion of nutrients and to aid in waste removal

[62�67]. These so-called liver-on-chip platforms allow for

the study of biochemical and mechanical cues such as

growth factor gradients and shear stress. Lee et al. demon-

strated that the presence of human stellate cells and applica-

tion of shear via flow enabled the formation of stable hepatic

spheroids on chip [68]. Furthermore, by linking microfluidic

channels between multiple tissue models, multiorgan phe-

nomena such as drug metabolite toxicity and disease pro-

gression can be captured in vitro, which is not possible in

traditional cultures [69�71].

Controlling three-dimensional architecture and

cellular organization

Another approach to improving the functionality of

tissue-engineered constructs is to more closely mimic

in vivo microarchitecture by generating scaffolds with a

highly defined material and cellular architecture, which

would provide better control over the 3D environment at

the microscale.

A range of rapid prototyping and patterning strategies

have been developed for polymers using multiple modes

of assembly, including fabrication using heat, light, adhe-

sives, or molding, and these techniques have been exten-

sively reviewed elsewhere [72]. For example, 3D printing

with adhesives combined with particulate leaching has

been utilized to generate porous poly(lactic-co-glycolic

acid (PLGA) scaffolds for hepatocyte attachment [73],

and microstructured ceramic [74] and silicon scaffolds

[75,76] have been proposed as platforms for hepatocyte

culture. Furthermore, molding and microsyringe deposi-

tion have been demonstrated to be robust methods for fab-

ricating specified 3D PLGA structures toward the

integration into implantable systems [77].

Microfabrication techniques have similarly been

employed for the generation of patterned cellular hydrogel

constructs. For instance, microfluidic molding has been

used to form biological gels containing cells into various

patterns [78]. In addition, syringe deposition in conjunc-

tion with micropositioning was recently illustrated as a

means to generate patterned gelatin hydrogels containing

hepatocytes [79]. Patterning of synthetic hydrogel systems

has also recently been explored. Specifically, the photopo-

lymerization property of poly(ethylene glycol) (PEG)

hydrogels enables the adaptation of photolithographic

techniques to generate patterned hydrogel networks. In

this process, patterned masks printed on transparencies

act to localize the ultraviolet exposure of the prepolymer

solution and, thus, dictate the structure of the resultant

hydrogel. The major advantages of photolithography-

based techniques for patterning of hydrogel structures are

its simplicity and flexibility. Photopatterning has been

employed to surface pattern biological factors [80], pro-

duce hydrogel structures with a range of sizes and shapes

[81,82], as well as build multilayer cellular networks

[83,84]. Consequently, hydrogel photopatterning technol-

ogy is ideally suited for the regulation of scaffold archi-

tecture at the multiple length scales required for

implantable hepatocellular constructs. As a demonstration

of these capabilities, photopatterning of PEG hydrogels

was utilized to generate hepatocyte and fibroblast cocul-

ture hydrogels with a defined 3D branched network,

resulting in improved hepatocyte viability and functions

under perfusion [85]. More recently, a “bottom-up”

approach for fabricating multicellular tissue constructs

utilizing DNA-templated assembly of 3D cell-laden

hydrogel microtissues demonstrates robust patterning of

cellular hydrogel constructs containing numerous cell

types [86]. Also, the additional combination of photopat-

terning with dielectrophoresis-mediated cell patterning

enabled the construction of hepatocellular hydrogel

742 PART | ELEVEN Gastrointestinal system

structures organized at the cellular scale. Overall, the abil-

ity to dictate scaffold architecture coupled with other

advances in scaffold material properties, chemistries, and

the incorporation of bioactive elements (discussed further

in the “Extracellular matrix for cell therapies” section)

will serve as the foundation for the future development of

improved tissue-engineered liver constructs that can be

customized spatially, physically, and chemically.

In vivo models

While there have been impressive advances in cell culture

models of the human liver, experimental animal models

still play an important role in the effort to engineer liver

therapies. Commonly performed surgeries such as bile

duct ligation and partial hepatectomy are experimentally

tractable models of acute liver injury, yet they are of little

clinical relevance. Drug-induced (e.g., carbon tetrachlo-

ride, acetaminophen, or thioacetamide) hepatotoxicity to

induce necrotic lesions is more recapitulative of human

pathophysiology, but the phenotype is difficult to repro-

duce [87]. In addition, modeling chronic liver injury in

animal models is problematic because they tend to rapidly

correct severe hepatic damage after a few days, which is

not representative of human disease progression and reso-

lution [88].

In order to model a human-like context for liver inju-

ries, it is necessary to develop improved, controlled mod-

els of human liver injury. Such animal models can be

useful for the evaluation of human liver biology and the

preclinical performance of candidate therapies and drugs

[89,90]. To accomplish this, human hepatocytes can be

transplanted orthotopically in immunocompromised mice

with no liver injury via the injection of a cell solution and

are useful for modeling human-specific drug metabolism,

liver injury, and hepatotropic infections. However, on

average, HT exhibits poor levels of engraftment

(B10%�30%) [91].

Transplanted hepatocytes have the ability to expand

preferentially if the host is compromised by injury or

genetic modification. The first genetically engineered

mouse model to demonstrate this was the Alb-uPA mouse

that carries a uroplasminogen activator under an albumin

promoter, causing liver injury and failure [92]. Aiming to

improve on the Alb-uPA system, a transgenic model of

hereditary tyrosinemia I was developed, in which a

genetic knockout of FAH leads to the hepatotoxic accu-

mulation of fumarylacetoacetate [93]. FAH knockout

mouse injury initiation and duration can be controlled

through the administration of a small molecule drug

[2-(2-nitro-3-trifluoro-methylbenzoyl)-1,3-cyclohexane-

dione] in the drinking water. Another inducible model

called TK-NOG, which expresses thymidine kinase

under an albumin promoter, causes hepatocyte ablation

following activation by ganciclovir treatment [94]. In

the AFC8 injury model, induction by the small mole-

cule AP20187 drives caspase-8-initiated apoptosis of

hepatocytes modified to express FK506 under an albu-

min promoter [95]. For all of the above models,

engraftment rates surpassing 70% have been observed.

A classical study in parabiotic rats in the 1960s

revealed that hepatic injury results in the expression of

systemic, soluble signals that have the potential to drive

liver regeneration [96]. Despite decades of research, the

complex signaling cascade driving liver regeneration is

still not well understood but has found utility in ectopic

humanized mouse models. Chronic liver injury often pre-

sents with high portal pressures, which can reduce

engraftment levels during an orthotopic transplantation.

Thus transplantation to ectopic sites is clinically attrac-

tive. Ectopic grafts that anastomose to the host vascula-

ture can interact with regenerative stimuli from the host

liver, causing expansion and proliferation of the trans-

planted human hepatocytes. Initially, ectopic transplanta-

tion was demonstrated in the LN [40] and spleen [97],

and later in the subcutaneous space and mesenteric fat

pad, both of which offer ease of accessibility for manipu-

lation and noninvasive imaging [58,59,98].

Taken together, the range of liver injury animal mod-

els are an essential tool for studying various perturbations

to normal liver biology and building implantable tissue

constructs to address acute and chronic liver failure. The

field is just beginning to uncover mechanisms that control

liver regeneration in various disease and injury contexts.

The discovery of new soluble regenerative signals will be

central to advancing therapies that have the potential to

improve the supply of donor tissue.

Cell sourcing

Cell number requirements

The development of cell-based therapies poses myriad

challenges, partially stemming from the scale of the liver.

An adult human liver is estimated to possess 241 billion

hepatocytes, 24 billion stellate cells, and 96 billion

Kupffer cells [99]. Sourcing such enormous cell numbers

using current technologies is not feasible.

However, many human HT studies suggest that clini-

cal intervention is possible with a fewer number of cells

and offer critical insights to help us determine minimum

cell numbers. In a review, Fisher and Strom cataloged 78

different human HT studies, detailing both the input cell

number and a qualitative description of functionality [30].

Correlating these, we can broadly surmise that to correct

inborn errors of metabolism, at least 1�10 billion hepato-

cytes are needed. However, for ALF, that number grows

to 5�20 billion cells. For liver cirrhosis, HTs have largely

been unsuccessful (discussed further in the “Hepatocyte

transplantation” section); therefore it is unclear what the

Hepatic tissue engineering Chapter | 40 743

cellular requirements for cirrhosis are. While injection of

hepatocytes is not the same as implantation within a scaf-

fold, these studies serve as useful inputs into more com-

plex physiological models.

In order to get us closer to these numbers, many dif-

ferent cell sources have been explored.

Immortalized cell lines

Immortalized hepatocyte cell lines can be derived from

liver tumor tissue or directly from primary hepatocytes

in vitro. The prominent lines utilized today are HepG2,

derived from hepatocellular carcinoma; HepaRG [100], a

human bipotential progenitor cell line; C3A, derived from

HepG2s; and Huh7, derived from liver tumor [101].

Several other fetal and adult hepatic cell lines have also

been established, typically using a combination of viral

oncogenes and the human telomerase reverse transcriptase

protein [102]. However, these cell lines lack the full func-

tional capacity of primary adult hepatocytes and there is a

risk that oncogenic factors could be transmitted to the

patient, limiting their use as a cell source for transplanta-

tion therapies.

Primary cells

Unlike immortalized lines, PHHs can provide a whole

host of human liver-specific function. PHHs, however, are

limited in supply, and their phenotype is difficult to main-

tain in vitro. Many methods have been developed for

maintaining long-term functionality of hepatocytes

through the use of a variety of configurations and bioma-

terial constructs, which are further discussed in the “In

vitro models” section. Due to limitations in the supply of

mature human hepatocytes, many groups have attempted

to promote the expansion and proliferation of PHHs

in vitro. Peng et al. have shown that TNFα promotes the

expansion of hepatocytes in 3D cultures and enables serial

passaging and long-term culture for more than 6 months

[103]. In a similarly notable study, Hu et al. identified an

optimal cell culture cocktail consisting of B27 supplement

(without vitamin A), R-spondin, CHIR99021 (a Wnt ago-

nist), NAC, nicotinamide, gastrin, epidermal growth fac-

tor (EGF), TGFα, fibroblast growth factor (FGF)7,

FGF10, HGF, a TGFβ inhibitor (A83-01), and ROCK

inhibitor that led to long-term 3D organoid culture of

PHHs [104].

Fetal and adult progenitors

Given their ability to differentiate into diverse lineages

both in vitro and in vivo, iPSC and human embryonic

stem cell cultures can also be utilized to generate HLCs.

Various differentiation protocols have been applied to

these cultures to yield cell populations that exhibit some

phenotypic and functional characteristics of hepatocytes

[62,105�108]. These populations are termed HLCs

because of their expression of fetal proteins and fetal-like

cytochrome P450 profiles [109]. While they are distinct

from mature adult hepatocytes, HLCs can still serve as a

potential cell source in very specific contexts.

In addition to pluripotent cells, bipotential progenitor

cells can also serve as a source for hepatocytes. Huch

et al. delineated conditions that allow for long-term

expansion of adult bile duct-derived EpCAM1 bipoten-

tial progenitor cells from the human liver [110]. The

expanded cell population attained using their protocol is

stable at the chromosomal level and can be converted into

functional HLCs in vitro and in vivo [110].

Reprogrammed hepatocytes

HLCs can also be generated using direct reprogramming

of mature cell types. For example, several groups have

demonstrated the feasibility of reprogramming fibroblasts

into HLCs without a pluripotent intermediate [111�113].

Cheng et al. demonstrated that a combination of nuclear

factors can stimulate the conversion of hepatoma cells to

HLCs [101]. These findings raise the future possibility of

deriving human HLCs directly from another adult cell

type.

Extracellular matrix for cell therapies

The ECM of the liver provides a structural scaffold with

bioactive cues that modulate hepatic function and pro-

mote vascularization. Collagen and fibronectin are the

major structural components of the liver ECM. Along

with other nonstructural proteins, these components par-

ticipate in integrin-mediated signaling between cells and

their surrounding matrix. Hepatocytes are sensitive to

their ECM, and it has been demonstrated that the presence

of abnormal amounts and/or types of ECM components

correlates with the onset and progression of liver fibrosis

[114].

ECM scaffolds for hepatic tissue engineering are use-

ful for constructing 3D tissue models and as a delivery

vehicle for implants. Polymeric biomaterial hydrogels

gained popularity as an engineering tool for recapitulating

a physiologically relevant 3D tissue niche. Aside from

creating a permissive environment for hepatocyte survival

and growth, ECM scaffolds for hepatic tissue engineering

also enable the formation of biliary and vascular networks

that will be further discussed in the “Vascular and biliary

tissue engineering” section. Broadly speaking, ECM scaf-

folds can be constructed using synthetic and/or naturally

derived polymers.

744 PART | ELEVEN Gastrointestinal system

Natural scaffold chemistry and modifications

A wide range of natural biomaterial polymers spanning

polysaccharides (e.g., dextran and chitosan), peptides

(e.g., collagen and fibrin), decellularized ECM (dECM),

and composites of these have been employed as hepatic

tissue�engineering scaffolds [55,57,58,115�121]. The

advantages of biologically derived materials include their

biocompatibility; naturally occurring cell adhesive moie-

ties; and, in the case of decellularization, native architec-

tural presentation of ECM molecules. However, naturally

derived biomaterials have several barriers to use in the

clinic, primarily due to lot-to-lot variability and xenoge-

neic origin.

The choice of material determines the physicochemi-

cal and biological properties of the scaffold. For

example, early efforts in developing implantable hepatic

constructs utilized collagen-coated dextran microcarriers

that enabled hepatocyte attachment since hepatocytes

are known to be anchorage-dependent cells. The intra-

peritoneal transplantation of these hepatocyte-attached

microcarriers resulted in successful replacement of

liver functions in two different rodent models of genetic

liver disorders [122]. Subsequently, collagen-coated or

peptide-modified cellulose [120,123], gelatin [124], and

gelatin�chitosan composite [125] microcarrier chemis-

tries have also been explored for their capacity to promote

hepatocyte attachment. On the other hand, materials that

are poorly cell adhesive such as alginate [115] have been

exploited for their utility in promoting hepatocy-

te�hepatocyte aggregation (i.e., spheroid formation) and

phenotypic stabilization within these scaffolds.

Collectively, the size of engineered tissues created by

these approaches is limited by oxygen and nutrient diffu-

sion to only a few hundred microns in thickness.

To address this constraint, recent work has sought to

use decellularized whole organ tissue as a matrix for liver

tissue engineering. The decellularization process utilizes

perfusion-based technologies to remove cells from donor

tissues but preserve the structural and functional charac-

teristics of the native underlying tissue. Recent advances

in decellularization protocols have yielded scaffolds with

native liver ECM composition, growth factor presentation,

vascular structure, and biliary network architecture

[126�128]. To date, seeding protocols have achieved up

to 95% efficiency of recellularization with relevant cell

populations (e.g., hepatocytes, vascular cells and bipotent

hepatic progenitors); resulting recellularized grafts exhib-

ited liver-specific function, and survival after transplanta-

tion in rodents [121,126,129]. Furthermore, cell-laden,

xeno-derived dECM scaffolds are compatible with immu-

nocompetent animal models [128]. However, given the

shortage of donor tissue, the wide use of dECM scaffolds

is unlikely.

Synthetic scaffold chemistry

In contrast to biologically derived material systems, syn-

thetic materials enable precisely customized architecture

(porosity and topography), mechanical and chemical

properties, and degradation modality and kinetics, which

are known to drive cell behavior. Synthetic materials that

have been explored for liver tissue engineering include

poly(L-lactic acid) (PLLA), PLGA, poly(ε-caprolactone),and PEG [85,98,130�135]. Polyesters such as PLLA and

PLGA are the most common synthetic polymers utilized

in the generation of porous tissue-engineering constructs.

These materials are biocompatible, biodegradable, and

have been used as scaffolds for HT [132,136]. A key

advantage of PLGA is the potential to finely tune its deg-

radation time due to differences in susceptibility to hydro-

lysis of the ester groups of its monomeric components

(lactic acid and glycolic acid). However, the accumulation

of hydrolytic degradation products has been shown to pro-

duce an acidic environment within the scaffold which

initiates peptide degradation and stimulates inflammation,

which may affect hepatocyte function [137].

Consequently, as alternatives to macroporous scaffold

systems, approaches aimed at the efficient and homoge-

neous encapsulation of hepatocytes within a fully 3D

structure have been explored. In particular, hydrogels that

exhibit high water content and thus similar mechanical

properties to tissues are widely utilized for various tissue-

engineering applications, including hepatocellular plat-

forms. Synthetic, PEG-based hydrogels have been

increasingly utilized in liver tissue-engineering applica-

tions due to their high water content, hydrophilicity, resis-

tance to protein adsorption, biocompatibility, ease of

chemical modification, and the ability to be polymerized

in the presence of cells, thereby enabling the fabrication

of 3D networks with uniform cellular distribution [138].

PEG-based hydrogels have been used for the encapsula-

tion of diverse cell types, including immortalized and pri-

mary hepatocytes and hepatoblastoma cell lines

[85,98,135]. The encapsulation of primary hepatocytes

requires distinct material modifications [e.g., 10% w/v

PEG hydrogel, inclusion of RGD adhesive motifs] as

detailed below, as well as, analogous to 2D coculture sys-

tems, the inclusion of nonparenchymal supporting cell

types such as fibroblasts and endothelial cells [135].

Modifications in scaffold chemistry

The relatively inert nature of synthetic scaffolds allows

for the controlled incorporation of chemical/polymer moi-

eties or biologically active factors to regulate different

aspects of cellular function. Chemical modifications such

as oxygen plasma treatment or alkali hydrolysis of PLGA

[139,140], or the incorporation of polymers such as

Hepatic tissue engineering Chapter | 40 745

poly(vinyl alcohol) or poly(N-p-vinylbenzyl-4-O-β-D-galactopyranosyl-D-glucoamide) (PVLA) into PLGA or

PLLA scaffolds [132,141,142] have improved hepatocyte

adhesion by modulating the hydrophilicity of the scaffold

surface [143]. Biological factors may include whole bio-

molecules or short bioactive peptides. Whole biomole-

cules are typically incorporated by nonspecific adsorption

of ECM molecules such as collagen, laminin, or fibronec-

tin [139,144]; covalent conjugation of sugar molecules

such as heparin [145,146], galactose [131,147], lactose

[145] or fructose [148]; or growth factors such as EGF

[149]. Alternatively, short bioadhesive peptides that inter-

act with cell surface integrin receptors have been exten-

sively utilized to promote hepatocyte attachment in

synthetic scaffolds. For example, conjugation of the RGD

peptide to PLLA has been shown to enhance hepatocyte

attachment [150], whereas RGD modification signifi-

cantly improved the stability of long-term hepatocyte

function in PEG hydrogels [98,135]. The additional incor-

poration of adhesive peptides that bind other integrins

may serve as a way to further modulate and enhance

hepatocyte function within synthetic polymer substrates.

Moreover, Stevens et al. demonstrated that integration of

matrix metalloproteinase-sensitive peptide sequences into

hydrogel networks as degradable linkages has been shown

to enable cell-mediated remodeling of the hydrogel [151].

The capacity to modify biomaterial scaffold chemistry

through the introduction of biologically active factors will

likely enable the finely tuned regulation of cell function

and interactions with host tissues important for

implantable systems.

Porosity

A common feature of many implantable tissue-

engineering approaches is the use of porous scaffolds that

provide mechanical support, often in conjunction with

cues for growth and morphogenesis. Collagen sponges,

various alginate and chitosan composites, and PLGA are

the most commonly used porous scaffolds for hepatocyte

culture and are generally synthesized using freeze-dry or

gas-foaming techniques. Pore size has been found to regu-

late cell spreading and cell�cell interactions, both of

which can influence hepatocyte functions [116], and may

also influence angiogenesis and tissue ingrowth [152].

Porous, acellular scaffolds are normally seeded using

gravity or centrifugal forces, capillary action, convective

flow, or through cellular recruitment with chemokines,

but hepatocyte seeding is generally heterogeneous in these

scaffolds [153,154].

Vascular and biliary tissue engineering

Beyond compatibility with hepatic cell types, scaffolds

should also be conducive to vasculature formation.

Relying on vascularization by the host is not sufficient for

large tissue constructs required for the clinic, because

cells that are not near capillary structures

(. 150�200 μm) are at a risk for necrosis after a matter

of hours due to a lack of oxygen, nutrient availability, and

waste transport. In this section, we discuss composite

approaches toward building scalable, vascularized

constructs.

Vascular engineering

Approaches to engineering vessels can generally be cate-

gorized as bottom-up induction of vascular assembly and

top-down fabrication of vascular conduits [155]. Bottom-

up vascular engineering approaches are built upon the

idea of neovascularization, or new vessel formation.

Vessel formation can occur by angiogenic sprouting, the

formation of vessels branching off of an existing blood

vessel, or vasculogenesis, the self-assembly of single

endothelial cells or progenitor cells into lumenized ves-

sels. Despite the ability of single vascular cells to coa-

lesce to enable self-assembly, vasculogenesis is

accelerated by coculture with supporting stromal cells,

such as fibroblasts, mesenchymal stem cells, and pericytes

[155�157]. Studies exploring angiogenesis and sprouting

events suggest that chemical gradients, fluid-driven shear

stress, and a hypoxia are key players in vessel formation

[158�161].

Top-down fabrication approaches dictate geometry

and architecture, rather than driving self-assembly.

Polymer molding using microetched silicon has been

shown to generate extensive channel networks with capil-

lary dimensions, though it is not amenable to high-

throughput manufacturing [162]. While direct printing of

cells can be cytotoxic, 3D printing has become a popular

approach for fabricating hollow channels that enable vas-

cular cell seeding. The challenge lies in using this

approach to build patent capillary beds, which are

5�10 μm in diameter. While two-photon polymerization

has an impressively high feature size resolution at

100 nm, the trade-off between build volume (i.e., building

constructs large enough for clinical impact), build speed

(i.e., impacting manufacturability), and printer resolution

renders it inappropriate for most applications in tissue

engineering. Alternate printers using direct-ink writing

(DIW) of viscoelastic materials have emerged as a power-

ful tool for the fabrication of patterned hydrogel con-

structs. DIW can achieve minimum feature size

resolutions from 1 to 250 μm, depending on the ink

“building block” size [163]. However, DIW printing

requires yield-stress fluid inks with restrictive viscosities

(102�106 mPa s), such that they fluidize under stress but

regain the original shear elastic modulus after printing

[163]. Kolesky et al. demonstrated that DIW and fugitive

inks were useful for multimaterial printing as well as

746 PART | ELEVEN Gastrointestinal system

construction of thick (. 1 cm), vascularized tissues

[164,165]. Miller et al. demonstrate the use of thermal

microextrusion approach to create sacrificial sugar glass

lattices that can be embedded in various cell-laden bioma-

terials, evacuated and subsequently lined with vascular

cells [166].

Taken together, self-assembly driven formation of a

capillary bed can be combined with printing of larger ves-

sels to enable the fabrication of a fully vascularized tissue

construct. Song et al. used 3D printing to create curved

vascular channels using a fugitive ink [159]. The vascular

cells lining the channel were able to degrade the sur-

rounding hydrogel matrix and undergo sprouting when

exposed to angiogenic factors. They further demonstrated

that the curvature and complexity of printed vasculature

impacted the extent of sprouting, which will be an impor-

tant consideration for vessel patterning in future clinical

applications [159].

Host integration

For vascularized constructs to survive after implantation

the engineered vessels must anastomose with the host vas-

culature. To date, the exact mechanism that drives inte-

gration with the host is not well understood [155].

A number of studies have elucidated the contribution of

cytokines important in angiogenesis and recruitment of

host vasculature in implant constructs, such as vascular

EGF (VEGF) [136], basic FGF [167], and VEGF in com-

bination with platelet-derived growth factor [168].

Furthermore, preimplantation of VEGF releasing alginate

scaffolds prior to hepatocyte seeding was demonstrated to

enhance capillary density and improve engraftment [169].

In addition, while building an AL tissue construct,

Stevens et al. demonstrated that the coimplantation of par-

allel endothelial cell cords with PHH and fibroblast spher-

oids led to better hepatic performance and survival when

compared to an implant with nonpatterned endothelial

cells, suggesting that there may be an optimal templated

endothelial geometry that enables vascularization and

host integration in vivo [58]. Surgical anastomosis poses

an alternative to biologically driven anastomosis, though

this approach requires invasive surgery and access to

suture-able vessels both in the graft and in the host.

Strategies to incorporate vasculature into engineered con-

structs include the microfabrication of vascular units with

accompanying surgical anastomosis during implantation

[75,170].

In addition to interactions with the vasculature, inte-

gration with other aspects of host tissue will constitute

important future design parameters. For instance, incorpo-

ration of hydrolytic or protease-sensitive domains into

hepatocellular hydrogel constructs could enable the degra-

dation of these systems following implantation [151]. Of

note, liver regeneration proceeds in conjunction with a

distinctive array of remodeling processes such as protease

expression and ECM deposition. Interfacing with these

features could provide a mechanism for the efficient inte-

gration of implantable constructs. Similarly to whole liver

or cell transplantation, the host immune response follow-

ing the transplant of tissue-engineered constructs is also a

major consideration. Immunosuppressive treatments will

likely play an important role in initial therapies, although

stem cell�based approaches hold the promise of

implantable systems with autologous cells. Furthermore,

harnessing the liver’s unique ability to induce antigen-

specific tolerance [171] could potentially represent

another means for improving the acceptance of engi-

neered grafts.

Biliary network engineering

Importantly, future iterations of hepatic grafts should

include a biliary system, which is responsible for excre-

tory functions. In a similar vein, we envision that a com-

bination of “top-down” manufacturing, such as the

aforementioned technologies for generating patent vascu-

lar conduits, and “bottom-up” approaches, which could

involve leveraging biological phenomena that drive bili-

ary morphogenesis, will be useful for building a biliary

network.

The biliary tree is a complex, 3D network of tubular

conduits of various sizes and properties. The liver con-

tains an intrahepatic compartment that is lined by biliary

epithelial cells, termed cholangiocytes, that aid in the

modification and removal of hepatocyte-secreted bile.

Even though the blood vessel�fabrication approaches

delineated above have not yet been applied to engineering

bile networks for implantable liver constructs, advance-

ments in cholangiocyte sourcing methods have made it

feasible. Sampaziotis et al. identified a protocol for

directed differentiation of human iPSCs into

cholangiocyte-like cells [172]. In 2017 the same group

provided the first proof-of-concept study to reconstruct

the gallbladder wall and repair the biliary epithelium

using human primary cholangiocytes expanded in vitro

[173]. Furthermore, current studies are focused on the

development of in vitro models that exhibits biliary mor-

phogenesis and recapitulates appropriate polarization and

bile canaliculi organization [174�176], as well as plat-

forms for the engineering of artificial bile duct structures

[177,178].

Conclusion and outlook

Traditionally, hepatic tissue�engineering research has

focused on designing the microenvironment to support a

stable hepatic phenotype. As concomitant advances in

pluripotent cell research and polymer chemistry have

been actualized, new cell sources and extracellular

Hepatic tissue engineering Chapter | 40 747

matrices have been added to the pipeline. This interdisci-

plinary synergy has been the driving force behind the

development of tissue-engineered grafts with long-term

survival and growth. In order to inch closer to the regen-

erative medicine north star of an ex vivo engineered graft

that can serve as a replacement for the native organ, there

are several areas to consider for improvement: (1) vascu-

larization of thick, dense grafts through a combination of

self-assembly and bioprinting; (2) engineering of the

hepatic graft to prevent immune rejection in allogeneic

and xenogeneic settings; (3) improved understanding of

metabolic and cellular requirements of various liver dis-

eases; (4) development of scalable, renewable cell sources

that do not compromise the functional capabilities of

cells; (5) leveraging of animal injury models as bioreac-

tors for cell sourcing; and (6) upscaling of grafts in a

manner that is compatible with FDA’s Good

Manufacturing Practice standards.

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